Integrated active flux microfluidic devices and methods

ABSTRACT

The invention relates to a microfabricated device for the rapid detection of DNA, proteins or other molecules associated with a particular disease. The devices and methods of the invention can be used for the simultaneous diagnosis of multiple diseases by detecting molecules (e.g. amounts of molecules), such as polynucleotides (e.g., DNA) or proteins (e.g., antibodies), by measuring the signal of a detectable reporter associated with hybridized polynucleotides or antigen/antibody complex. In the microfabricated device according to the invention, detection of the presence of molecules (i.e. Polynucleotides, proteins, or antigen/antibody complexes) are correlated to a hybridization signal from an optically-detectable (e.g. fluorescent) reporter associated with the bound molecules. These hybridization signals can be detected by any suitable means, for example optical, and can be stored for example in a computer as a representation of the presence of a particular gene.

PRIORITY APPLICATIONS

This is a continuation-in-part of copending U.S. patent application Ser.No. 09/724,548 filed on Nov. 28, 2000. U.S. patent application Ser. No.09/724,548 claims priority under 35 U.S.C. §119(e) to copending U.S.provisional patent application Ser. Nos. 60/209,243; 60/211,309; and60/249,360 filed on Jun. 5, 2000; Jun. 13, 2000; and Nov. 16, 2000,respectively. Each of these priority applications hereby incorporated,by reference, in its entirety.

The present invention was made with Government support under Grant No. 5R29HG01642-03 awarded by the National Institutes of Health. The UnitedStates Government may have certain rights to this invention pursuant tothese grants.

1. FIELD OF THE INVENTION

This invention relates to microfluidic devices and methods, includingmicrofabricated multilayer elastomeric devices with active pumps andvalves. More particularly, the devices and methods of the inventioncomprise a loop channel that is selectively open or closed to at leastone input or output, and which actively circulates a fluid received inthe loop. The loop can be closed by microvalves, for example elastomericmicrovalves interposed between an inlet or outlet channel and the loopchannel. Any fluid, such as a liquid (preferably aqueous), gas, slurry,etc. can be moved through fluid channels of the microfluidic device,which are typically on an elastomeric fluid layer and comprise the loopchannel and its inlet and outlet channel or channels. Fluid within theloop is circulated, for example by active pumping, which can be donewhile the loop is open or closed to any or all channels that communicatewith the loop channel. Pumping can be provided by a series of at leastthree microvalves which cooperate to form a peristaltic pump by cyclingthrough an appropriate sequence of on/off or open/close steps.

Microvalves are formed and actuated by control lines or channels,typically on an elastomeric control layer adjacent to a fluid layer. Amicrovalve is formed by the elastomeric interchannel membrane separatinga fluid channel on one layer and an appropriately placed control line onan adjacent layer, where the fluid channels and control lines cross.Fluid in a control line, preferably a pressurized gas and mostpreferably air, can selectively deform or release the interchannelmembrane of a microvalve, to close or open the valve and restrict orpermit flow in the adjacent cooperating fluid channel.

The loop channel can be provided with any reagents or reactants to bemixed or combined for any purpose, including any chemical reactions orinteractions. In one embodiment, molecules are applied to a surface thatis exposed to fluid circulating in the loop, to facilitate a desiredinteraction between the molecules and one or more components of thefluid. For example, DNA probes can be patterned onto spots in the loopchannel for analysis of a DNA sample, by analyzing (e.g. imaging) anyhybridization of probe DNA with sample DNA.

Thus, the devices and methods comprise integrated diagnostic chips withelastomeric channels, surface patterning, and surface chemistriesadapted for multiparameter analysis of a sample, e.g. DNA hybridization.Flow control, reagent metering, in-line mixing, loop circulations, and“rotary” designs are also described. These devices can be used for“lab-on-a-chip” applications, for example to test for and diagnosemultiple diseases. Devices and methods include detection of organisms orgenetic disorders, or determining a genetic predisposition orsusceptibility of humans and animals to genetic disorders, cancer andcancer-related diseases. Microfabricated chips of the invention can beused to measure gene expression, to detect the presence of pathogenicorganisms or DNA, for DNA fingerprinting and forensic analysis, and forother applications in which molecules, viruses, particles, or cells andthe like are analyzed, identified, evaluated, tested or sorted.

The invention also relates to methods for the rapid diagnosis of diseaseby detecting molecules (e.g. amounts of molecules), such aspolynucleotides (e.g., DNA) or proteins (e.g., antibodies), by measuringthe signal of a detectable reporter associated with the molecules (e.g.,fluorescent, ultraviolet, radioactive, color change, or another signal).Preferably, the reporter or its signal is optically detectable. In theseembodiments, a positive result (i.e. the presence or absence of theparticular gene or antigen) is correlated to a signal from anoptically-detectable reporter associated with hybridized polynucleotideor antigen/antibody complex. These polynucleotides or complexes can alsobe identified, assessed, or sorted (e.g. by size) in a microfabricateddevice that analyzes the polynucleotides according predeterminedalgorithms or characteristics, for example restriction fragment lengthpolymorphism (RFLP).

Certain embodiments of the invention comprise an integrated microfluidicsystem with an array of diagnostic probes attached to a substrate.Multiple disease diagnosis of a sample can be done by using DNAhybridization, antibody/antigen reaction, or other detection methods.The loaded sample is actively moved in a loop on the device by abuilt-in peristaltic pump. Target DNA or antibodies in the sample, ifany, associate or bind with their matching probes and give a positivesignal of the corresponding diseases. The invention provides enhancedhybridization rates and improved speed and efficiency by active pumping,(e.g. ˜20 minutes for 30 probes). The devices and methods of theinvention are also accurate and require very little amount of sample,e.g. only a few microliters of total volume and a few target DNAmolecules or antibodies for each disease; e.g. less than 100, preferablyless than 50 molecules. The system is also advantageously small,typically 1 inch by 0.1 inch, and is easy and inexpensive to fabricate.It is disposable and thus eliminates cross-contamination. Many samplepreparation and/or treatment steps can be incorporated into the device.

Other advantages include that multiple diseases can be diagnosedrapidly, contemporaneously or simultaneously on a single chip, e.g. inserial or in parallel, making disease diagnosis simpler and less costly.Automation can also be used. Another advantage is that there is no needto custom-design each chip for each application: the invention is highlyflexible in design and use. Additional functions can be incorporated asdesired, such as in-line digestion, separation i.e., for DNAfingerprinting or RFLP analysis and other techniques such as insitu-enzymatic labeling, PCR, etc. Small samples can be processedquickly, easily and accurately without the need for PCR, and thuswithout the potential costs, complications, errors or otherdisadvantages of PCR.

2. BACKGROUND OF THE INVENTION

Diagnosis of the sources, types and cures of diseases is usually done bydoctors, based on symptoms and on simple tests and observations. Becausethere are so many similar diseases, further diagnoses are often requiredto precisely differentiate them, especially for diseases with infectiousor genetic roots, such as HIV, tuberculosis, hepatitis and human BRCA1breast cancer. Conventionally, disease diagnosis has been carried out bytechniques such as bacterial culture or antibody/antigen reactions (1).Recently, molecular techniques such as DNA restriction fragment lengthpolymorphism analysis (RFLP) have become more widely used for thedetection of mutation-intense diseases or for genotyping specificpathogenic microorganisms, e.g. tuberculosis (80). However, relativelylarge sample volumes have been necessary and significant manipulation ofthe sample may be required. The conventional techniques are costly, timeconsuming and very labor-intensive. These methods may not work when onlysmall samples are available. Rapid, contemporaneous, or simultaneoustesting for more than one organism, disease characteristic, or parametermay be impractical or impossible.

DNA chips have been developed for disease diagnosis, using an array ofvarious DNA hybridization probes laid down onto a solid substrate (2-4,72-76, 81-83). The probes in these techniques are designed to react onlywith specific target DNA fragments from chosen disease entities.Nevertheless, hundreds of microliters to a few milliliters of sample arerequired to cover the chip. A further drawback is that is that thediffusion constant of DNA fragments is small, on the order of ˜10⁻⁷cm²/sec for 1-kbp DNA fragments (5). Thus, passive diffusion is anextremely slow process for large molecules such as DNA. Diffusion ratescan be calculated using the equation:

l=√{square root over (Dt)},

where l is diffusion length, D is the diffusion constant and t is time.If D is 10⁻⁷ cm²/s for a typical 1 kbp DNA and t is one hour (3600seconds), the diffusion length/is 0.19 mm. It follows that for passivediffusion of the DNA, each hybridization spot can only cover an area ofabout 0.4 mm in diameter. Even after one day i.e., 24 hours, only targetsamples in an area of ˜2 mm in diameter can reach a specific probe togive a positive signal. Therefore, it takes a relatively long time fortarget DNA to be directed to complementary DNA probes. DNA may be lostor fail to find a matching probe, or will not do so in a reasonabletime. PCR amplification may be needed to obtain enough DNA sample, whichcomplicates the process and gives new sources of possible errors.

The invention addresses these and other problems. Microfluidic chipshaving elastomeric channels are provided, and an active flow of sampleis delivered, for example by actively transporting a DNA or proteinsample around a central loop within the device by a built-in (on-chip)peristaltic pump. The pumping action improves the efficiency ofhybridization by directing the biological sample to it's target, whichobviates the need for larger sample volumes and avoids the longerreaction times needed for passive devices (e.g. sample diffusion). Amicrofabricated or microfluidic device may be used to implement thesetechniques, for example to detect or separate labeled fragments.Microfluidic devices and related techniques have been described (11, 2575-77, 84). These devices permit the manipulation, automatically ifdesired, of small volumes of biological samples on a small device, wherereactions and diagnoses may be carried out.

The invention also encompasses the identification and separation ofnucleic acid fragments by size, such as in sequencing of DNA or RNA.This is a widely used technique in many fields, including molecularbiology, biotechnology, and medical diagnostics. The most frequentlyused conventional method for such separation is gel electrophoresis, inwhich different sized charged molecules are separated by their differentrates of movement through a stationary gel under the influence of anelectric current. Gel electrophoresis presents several disadvantages,however. The process can be time consuming, and resolution is typicallyabout 10%. Efficiency and resolution decrease as the size of fragmentsincreases; molecules larger than 40,000 base pairs are difficult toprocess, and those larger than 10 million base pairs cannot bedistinguished.

Methods have been proposed for determination of the size of nucleic acidmolecules based on the level of fluorescence emitted from moleculestreated with a fluorescent dye. See Keller, et al., 1995 (42); Goodwin,et al., 1.993 (39); Castro, et. al., 1993 (38); and Quake, et al., 1999(70). Castro (38) describes the detection of individual molecules insamples containing either uniformly sized (48 Kbp) DNA molecules or apredetermined 1:1 ratio of molecules of two different sizes (48 Kbp and24 Kbp). A resolution of approximately 12-15% was achieved between thesetwo sizes. There is no discussion of sorting or isolating thedifferently sized molecules.

In order to provide a small diameter sample stream, Castro (38) uses a“sheath flow” technique wherein a sheath fluid hydrodynamically focusesthe sample stream from 100 μm to 20 μm. This method requires that theradiation exciting the dye molecules, and the emitted fluorescence, musttraverse the sheath fluid, leading to poor light collection efficiencyand resolution problems caused by lack of uniformity. Specifically, thismethod results in a relatively poor signal-to-noise ratio of thecollected fluorescence, leading to inaccuracies in the sizing of the DNAmolecules.

Goodwin (39) mentions the sorting of fluorescently stained DNA moleculesby flow cytometry. This method, employs costly and cumbersome equipment,and requires atomization of the nucleic acid solution into droplets,where each droplet contains at most one analyte molecule. Furthermore,the flow velocities required for successful sorting of DNA fragmentswere determined to be considerably slower than used in conventional flowcytometry, so the method would require adaptations to conventionalequipment. Sorting a usable amount (e.g., 100 ng) of DNA using suchequipment would take weeks, if not months, for a single run, and wouldgenerate inordinately large volumes of DNA solution requiring additionalconcentration and/or precipitation steps.

Quake (70) relates to a single molecule sizing microfabricated device(SMS) for sorting polynucleotides or particles by size, charge or otheridentifying characteristics, for example, characteristics that can beoptically detected. The invention includes a fluorescence activatedsorter (FAS), and methods for analyzing and sorting polynucleotides bymeasuring a signal produced by an optically-detectable (e.g.,fluorescent, ultraviolet or color change) reporter associated with themolecules. These methods and microfabricated devices allow for highsensitivity, no cross-contamination, and lower cost than conventionalgel techniques. In one embodiment of the invention, it has beendiscovered that devices of this kind can be advantageously designed foruse in molecular fingerprinting applications, such as DNAfingerprinting.

These and other devices, including those which provide single moleculeprocessing, can be used in combination with the loop channel andperistaltic pump devices of the invention. Likewise, other mechanisms offlow control, such as electroosmotics and electrophoresis, may be usedin addition to or in combination with the loop channel, pump and valvearrangements described herein.

Given the current state of the art, it is desirable to provide newdevices and methods for the rapid diagnosis of multiple diseases, e.g.by detecting the presence or absence of a particular gene. Such devicesand methods may include analyzing and sorting differently sized nucleicacid or protein molecules with high resolution. It is likewise desirableto provide microfluidic chip designs having an architecture suitable formultiparameter analysis, including for example the rapid,contemporaneous or simultaneous evaluation of a sample in a battery oftests, for a plurality of characteristics, or against an array oftargets or potential targets, for example by circulating sample in aloop channel for repeated exposure to a set of diagnostic probes.

3. SUMMARY OF THE INVENTION

The invention provides microfabricated devices and methods for the rapiddetection of DNA, proteins, viruses or other molecules or particles,e.g. associated with a particular disease. The device includes a chiphaving a microfabricated analysis unit, preferably microfabricated in oronto a substrate of the chip. Each analysis unit includes a main channelin communication with a sample inlet channel, a target (e.g.hybridization) loop, and a detection region. The target loop ispatterned with target molecules (e.g. polynucleotides or polypeptides).Additional channels may intersect or communicate with the target loop,on the same layer or on a different layer of the chip. Multilayerintegrated or monolithic devices are preferred. The detection region maycoincide with all or part of the target loop. The inlet channel maycomprise a plurality of channels communicating with each other or withone or more reservoirs, or with one or more feed channels, to controlflow or to deliver a plurality of reagents or samples. Typical devicesalso have an outlet channel, which may lead to an outlet reservoir. In apreferred embodiment, the target loop cooperates with a peristaltic pumpassembly. Adjacent and downstream from the detection region, the mainchannel may have a discrimination region or branch point leading to atleast two branch channels. In embodiments having an outlet channel, anoutlet channel may be placed anywhere on the chip, but typicallycommunicates with a main channel downstream of the detection region.Each channel may carry any fluid flow, e.g. a liquid (preferably anaqueous solution), a gas (preferably air), or a slurry.

Embodiments of these microfluidic devices are also provided whichcomprise a plurality of target loops, each of which is driven by a pumpsuch as a peristaltic pump. The plurality of target loops in thesedevices may also be interconnected by microfluidic channels. Forexample, in one embodiment each target loop is connected to a commonsample inlet and/or a common sample outlet by a common inlet or outletchannel, respectively. The inlet and/or outlet channels may, forexample, be fluidly connected to a plurality of branch channels, witheach branch channel connecting, in turn, to a particular target loop ofthe device. Alternatively, the plurality of target loops in thesemicrofluidic devices may be connected to separate sample inlets and/orsample outlets, e.g., by a separate inlet or outlet channel.

In a preferred multilayer device, a pattern of fluid channels isfabricated on one layer, and a pattern of air channels is fabricated ona second layer. In operation, the fluid channels of the device carry anyfluid, typically a liquid and most typically water or an aqueoussolution or slurry. These channels are typically used to receive,process, analyze and work with samples and reagents, and may also bereferred to as treatment channels. Air channels typically operate onanother layer of the device and may intersect or communicate with fluidchannels where adjacent layers of the device meet, for example atjunctions or at the interface of two adjacent layers. The air channelsmay carry any pressurized flow of any fluid, liquid or gas, although airis generally preferred. The air channels are typically used to controlthe flow of fluid in the fluid or treatment channels, for example usingair pressure, or by controlling microfabricated pumps and/or valvesintegrated on the chip. These channels can also be called controlchannels or control lines.

In certain embodiments, any layer of the device may have any kind ofchannel, in any pattern, array or arrangement. Channels in a multilayerdevice may also be made to encompass or transverse more than one layer,communicate with more than one layer, or to cross from one layer toanother; for example by fabricating overlapping adjacent layers havingoverlapping channels which intersect or meet in any desiredconfiguration or plane. Adjacent channels or layers in a multilayer mayare not necessarily in contact with each other, and may be separated bygaps between layers or between channels. Openings may be made inchannels as desired, for communication with other channels or layers, orfor communication with a gap between layers, which for example may houseor comprise one or more reservoirs. Any desired pattern or array ofchannels and intercommunications among and between them can be made byfabricating and joining corresponding negative molds of siliconelastomer according to the techniques described herein. See also, Ungeret al. (6).

In preferred embodiments, fluid or treatment channels are not open ordirectly connected to air or control channels. That is, they areindependent channel systems that do not directly feed into each other;they are sealed from each other and their contents do mix. The treatmentand control channels interact with each other where they intersect toform a microvalve. When sufficient pressure, e.g. air pressure isapplied to an air channel, the elastomeric membrane between the controlchannel and the treatment channel is deformed where the channelsintersect. Sufficient pressure pinches, restricts or closes off the flowin the treatment channel, forming a closed microvalve. The valve isopened by releasing the pressure in the control channel. Thus, valvescan be positioned as desired throughout a microfluidic device, each ofwhich can be operated independently or in combination to control theprocessing and flow in the treatment channels.

These valves are actuated by moving a portion of the ceiling, roof orwall of a channels itself (i.e. a moving membrane). Valves and pumpsproduced by these techniques have a zero dead volume, and switchingvalves made by this technique have a dead volume approximately equal tothe active volume of the valve, for example about 100×100×10 μm=100 pl.Such dead volumes and areas consumed by the moving membrane areapproximately two orders of magnitude smaller than microvalvesdemonstrated to date. Experimentally, the response of such valves hasbeen almost perfectly linear over a large portion of its-range oftravel, with minimal hysteresis. Accordingly, the present valves areideally suited for microfluidic metering and fluid control. The linearnature of the valve response demonstrates that the individual valves arewell modeled as Hooke's Law springs. Furthermore, high pressures in theflow channel (i.e. back pressure) can be countered simply by increasingthe actuation pressure. Experimentally, the present inventors haveachieved valve closure at back pressures of 70 kPa, but higher pressuresare also contemplated.

A preferred silicon elastomer for treatment and control channels isGeneral Electric Silicon RTV 615, made by combining the components RTV615A and RTV 615B. Transparent elastomers are particularly preferred. Incertain embodiments, the treatment and control channels may be made indifferent molds, i.e. on different layers, using different elastomers.

Air pressure can be controlled for example using external (off-chip)three-way pneumatic valves such as model LHDA1211111H (Lee Company) tomanipulate the on/off states of each individual microvalve. Valves mayalso be fabricated to have different stiffnesses, tolerances orthresholds, or different switching pressures, so that different valveswill open and close at different pressures along one control channel.This may be determined, for example, by the elastomers used, by theshape and dimension's of the channels, by the distances or gaps betweenintersecting treatment and control channels, and by the thickness of themembrane between them.

Three microvalves in a series become a peristaltic pump when anappropriate on/off pumping sequence is applied. This causes successivewaves of contraction along the treatment channels which propels thecontents of the channel onward. For example, a flow of sample can berouted through the treatment channels as desired, by appropriatelymanipulating the valves to form a peristaltic pump that drives the fluidin the desired direction and through the desired channels. Valves canalso be used to open and close channels as desired, to control thepattern and timing of flow.

In preferred embodiments, treatment channels are microfabricated into atransparent layer of a microfluidic device that is bonded to a glass orsimilar transparent or optically suitable probe substrate or coverslip,particularly in regions corresponding to the detection region. Thisprovides access to the channel or channels for optical detection, forexample by a high numerical aperture (NA) microscope. In a preferredembodiment, selected regions of the probe substrate corresponding toselected treatment channels are patterned with target or probemolecules, such as DNA, polynucleotide, protein, or antibody probes. DNAprobes corresponding to a set of different diseases can be laid down ona target loop to form distinct hybridization spots. In this embodimentthe target loop and its corresponding probe pattern preferably has acircular path on the face of the chip and its glass substrate. Any pathof any shape can be used, although a path which can be selectively openand closed is preferred. For example, the path of a loop channel can berectangular or square. The detection region in this embodiment comprisesany or all of the target loop. That is, sample introduced to the chip,for example by capillary action, will enter the target loop, andmolecules in the sample can bind to their corresponding probes on theglass substrate, if any. Binding can be detected using any suitabletechnique, including fluorescence, as described herein.

To improve the speed and accuracy of detection, minimize the amount ofsample needed, and address diffusion issues, microvalves can be used todrive a peristaltic pumping action as described, which moves the samplearound and around the target loop for continuous and or repeatedexposure to the probes. The sample passes each probe several or manytimes, meaning that all sample molecules (e.g. DNA) will eventually andrelatively quickly find and bind (hybridize) with matching targets (e.g.polynucleotide probes) at the right hybridization, spots. Little or nosample is wasted, PCR amplification may not be needed, and heating(preferably intermittent) can be applied to denature falsehybridizations and obtain more accurate results in successive passesthrough the target loop.

An object of the present invention is the simultaneous diagnosis ofmultiple diseases by detecting molecules (e.g. amounts of molecules),such as polynucleotides (e.g., DNA) or proteins (e.g., antibodies), bymeasuring the signal of a detectable reporter associated with thehybridized polynucleotides or antigen/antibody complexes.

An additional object of the invention is to provide a kit for the rapiddiagnosis of disease.

A further object of the present invention is to provide algorithms fordetermining the existence of specific disease targets.

A still further object of the present invention is to determine theseverity of a particular disease, for example according to the signalintensity from hybridization of a sample and target.

Another object of the present invention is to determine thesusceptibility or predisposition of patients to a particular disease.

Yet another object of the present invention to provide methods formixing two or more different fluids (i.e., fluids comprising differentmolecules or particles). Accordingly, the invention provides for the useof a microfluidic device to mix two or more different fluids.

Still another object of the present invention is to provide methods forbinding a sample (e.g., of nucleic acids, polypeptides, cells, virionsor other molecules and/or particles) to a target (for example, to amolecular probe, such as a complementary nucleic acid or an antibodyprobe). Accordingly, the invention also provides for the use of amicrofluidic device to bind a sample to a target.

Additional objects of the invention include measuring gene expressionlevels; sequencing DNA; “fingerprinting” DNA sequences; measuringinteraction of proteins, etc. with DNA sequences of length n (e.g. withall oligonucleotides of size n); and mutation and/or single nucleotidepolymorphism (SNP) detection.

Other objectives will be apparent to persons of skill in the art.

In accomplishing these and other objectives, the invention provides a“lab-on-a-chip” device which utilizes several orders of magnitude lowersample volumes than conventional methods. For example, rather than usinglarge sample volumes, a few droplets are enough. This reduces the useand cost of reagents and may reduce the risks to patients. The activedesign of the device increases the speed of the detection processsignificantly. A multiple disease diagnosis can be complete in a fewminutes. Furthermore, the device is inexpensive and disposable, due inpart to the materials used and the easy fabrication process. Automaticcomputer control can be easily integrated by controlling the switchingof pneumatic valves via electronic driving circuits. Therefore, manuallabor and chances of errors are greatly reduced. The invention offersflexibility of design and fabrication with the capability for many otherfunctions.

In a preferred embodiment, the substrate of the device is planar, andcontains a microfluidic chip made from a silicone elastomer impressionof an etched silicon wafer according to replica methods insoft-lithography. See e.g. the devices and methods described in pendingU.S. application Ser. No. 08/932,774 filed Sep. 25, 1997; No. 60/108,894filed Nov. 17, 1998; No. 60/086,394 filed May 22, 1998; and No.09/325,667 filed May 21, 1999 (molecular analysis systems). Thesemethods and devices can further be used in combination with the methodsand devices described in pending U.S. application Ser. No. 60/141,503filed June 28, 199; No. 60/147,199 filed Aug. 3, 1999 and No.60/186,856, filed Mar. 3, 2000 entitled “Microfabricated ElastomericValve and Pump Systems”. Each of these references is hereby incorporatedby reference in its entirety.

In a preferred embodiment, the microfabricated device is used for theidentification of particular genes within the genome of pathogenicorganisms, genetic disorders or genetic predisposition or susceptibilityof humans or animals to cancer and cancer-related diseases.Microfabricated methods and devices are fast and require only smallamounts of material, yet provides a high sensitivity, accuracy andreliability. In another embodiment, the microfabricated device can beused for detecting or sorting nucleotide fragments in a fingerprintaccording to size.

Microfabricated Device. The device includes a chip having a substratewith at least one microfabricated analysis unit. Each analysis unitincludes a main channel, having a sample inlet, typically at one end,having along the length of the main channel a target or hybridizationloop and a detection region, and having, an outlet or a branch pointdiscrimination region adjacent and downstream of the detection region,leading to a waste channel or to a plurality of branch channels. In oneembodiment two or more branch channels originate at the discriminationregion and communicate with the main channel. The analysis unit alsoprovides a stream or flow of solution, preferably but not necessarilycontinuous, which contains sample molecules and passes through thedetection region. In certain embodiments the detection region comprisesone or more regions of a target loop, is coextensive with the targetloop, or comprises a region corresponding to each hybridization spot onthe target loop. Thus, a device of the invention can comprise aplurality of detection regions, or one detection region comprisingdiscrete test areas or hybridization spots, and detection can beserially, in parallel, or all at once. The presence, absence or level ofreporter from each molecule is measured as it passes within thedetection region. In a certain embodiments, on average only one moleculeoccupies one or more detection regions at a time. If desired, themolecule is directed to a selected branch channel based on the presence,absence or level of reporter. In other embodiments the molecule is heldin the detection region, temporarily or permanently, for example bybinding to a probe.

In a preferred embodiment, the substrate is planar, and contains amicrofluidic chip made from a silicone elastomer impression of an etchedsilicon wafer using replica methods in soft-lithography (23). In oneembodiment, the main channel meets branch channels to form a “T” (Tjunction) at a discrimination point. A Y-shaped junction, and othershapes and geometries may also be used. A detection region is typicallyupstream from the branch point. Molecules or cells are diverted into oneor another outlet channel based on a predetermined characteristic thatis evaluated as each molecule passes through the detection region. Thechannels are preferably sealed to contain the flow, for example byfixing a transparent coverslip, such as glass, over the chip, to coverthe channels while permitting optical examination of one or morechannels or regions, particularly the detection region. In a preferredembodiment the coverslip is pyrex, anodically bonded to the chip.Alternatively, the substrate may be an elastomer, which may proveadvantageous when higher back pressures are used.

Other devices such as electrophoresis chips may also be used. Exemplarydevices are described in U.S. Pat. Nos. 6,042,709; 5,965,001; 5,948,227;5,880,690; and 6,007,690.

Channel Dimensions. The channels in a multiparameter molecular analysisdevice are preferably between about 10 μm and about 200 μm in width,typically 50-100 μm, and most preferably about 100 μm. The channels arepreferably about 2-20 μm in depth for DNA or polynucleotide analysis,more typically about 10 μm. The detection region in preferredembodiments has a volume of between about 1 μl and about 1 nl. A typical1 Kbp DNA fragment takes about 10 seconds to diffuse 10 μm, e.g. fromthe top of a treatment channel to the hybridization probes fixed to thebottom of the channel (e.g. on a glass substrate). In a cell analysisdevice the channels are preferably between about 1 and 500 microns inwidth and between about 1 and 500 microns in depth, and the detectionregion has a volume of between about 1 fl and 100 nl. The channels maybe of any dimensions suitable to accommodate the largest dimension ofthe molecules, particle, viruses, cells or the like to be analyzed.

Manifolds. A device which contains a plurality of analysis units mayfurther include a plurality of manifolds, the number of such manifoldstypically being equal to the number of branch channels in one analysisunit, to facilitate collection of molecules from corresponding branchchannels of the different analysis units.

Flow of Molecules. In one embodiment, the molecules are directed orsorted by electroosmotic force. A pair of electrodes apply an electricfield or gradient across the discrimination region that is effective tomove the flow of molecules through the device. In a sorting embodimentthe electrodes can be switched to direct a particular molecule into aselected branch channel based on the amount of reporter signal detectedfrom that molecule.

In another embodiment, a flow of molecules is maintained through thedevice via a pump or pressure differential, and a valve structure can beused at the branch point effective to permit each molecule to enter onlyone selected branch channel. Alternatively, a valve can be placed in oneor more channels downstream of the branch point to allow or curtail flowthrough each channel. In a related embodiment, pressure can be adjustedat the outlet of each branch channel effective to allow or curtail flowthrough the channel. Pump and valve arrangements are preferred, such asthose disclosed in Ser. No. 60/186,856 filed Mar. 3, 2000 entitled“Microfabricated Elastomeric Valve and Pump Systems”.

Microvalves acting in concert to form a pump are preferred forcirculating a fluid in a closed loop of the invention. For example,three or more valves in series comprise a peristaltic pump when actuatedin an appropriate sequence. Electroosmotic and electrophoretic drivesmay be less suitable or inoperable in certain applications, for exampledue to issues of electrical charge.

In preferred polynucleotide sorting embodiments, the concentration ofpolynucleotides in the solution is between about 10 fM and about 1 nMand the detection region volume is between about 1 fl and about 1 pl.The molecules can be diverted, for example, by transient application ofan electric field effective to bias (i) a molecule having the selectedproperty, such as size (e.g., between about 100 bp and about 10 mb) toenter one branch channel, and (ii) a molecule not having the selectedproperty to enter another branch channel. Alternatively, molecules canbe directed into a selected channel, based on a detectable property, bytemporarily blocking the flow in other channels, such that thecontinuous stream of solution carries the molecule having the selectedproperty into the selected channel. Pumps and valves may also be used todivert flow, and carry molecules into one or another channels, andmechanical switches may also be used. These methods can also be used incombination, and likewise molecules can be diverted based on whetherthey have a selected property or size, or do not have that property orsize, or exceed or do not exceed a selected threshold measurement.

Optical Detection. Preferably the molecules are optically detectablewhen passing through the detection region. For example the molecules maybe labeled with a reporter, for example a fluorescent reporter. Theoptically detectable signal can be measured, and generally isproportional to or is a function of a characteristic of the molecules,such as size, molecular weight, or affinity for a predetermined probe. Afluorescent reporter, generating a quantitative optical signal can beused. Fluorescent reporters are known, and can be associated withmolecules such as polynucleotides using known techniques. Intercalatingdyes, incorporation of fluorescent-labeled single nucleotides, DNAbeacons or other well-established detection schemes can be used todetermine the final diagnostic results. Suitable fluorescentintercalating dyes include YOYO-1, TOTO-1 and PicoGreen from MolecularProbes, Eugene, Oreg.

In a preferred molecular fingerprinting embodiment, the reporter labelis a fluorescently-labeled single nucleotides, such as fluorescein-dNTP,rhodamine-dNTP, Cy3-dNTP, Cy5-dNTP, where dNTP represents dATP, dTTP,dUTP or dCTP. The reporter can also be chemically-modified singlenucleotides, such as biotin-dNTP. Alternatively, chemicals can be usedthat will react with an attached functional group such as biotin.

Sorting Molecules. In another aspect, the invention includes a method ofisolating molecules, polynucleotides, proteins, viruses, particles,beads, cells, etc., for example polynucleotides having a selected size.The method includes: a) flowing a stream of solution containingreporter-labeled polynucleotides through a channel comprising adetection region having a selected volume, where the concentration ofthe molecules in the solution is such that the molecules pass throughthe detection region one-by-one, c) determining the size of eachmolecule as it passes through the detection region by measuring thelevel of the reporter, in the stream, and (d) diverting (i) moleculeshaving the selected size into a first branch channel, and (ii) moleculesnot having the selected size into a second branch channel.Polynucleotides diverted into any channel can be collected as desired.

Synchronization. In each embodiment where molecules are diverted, asopposed to being measured only, the molecules are detected one-by-onewithin a detection region, and are diverted one-by-one into theappropriate channels, by coordinating or synchronizing the diversion offlow with the detection step and with the flow entering the detection,as described for example in more detail below. In certain embodimentsthe flow rate may be adjusted, for example delayed, to maintainefficient detection and switching, and as described below the flow mayin some cases be temporarily reversed to improve accuracy.

Sizing Molecules. In yet another aspect, the invention includes a methodof sizing polynucleotides in solution. This method includes: a) flowinga continuous stream of solution containing reporter-labeledpolynucleotides through a microfabricated channel comprising a detectionregion having a selected volume, where the concentration of themolecules in the solution is such that most molecules pass through thedetection region one by one, and b) determining the size of eachmolecule as it passes through the detection region by measuring thelevel of the reporter.

Multiparameter and High Throughput Embodiments. In addition to analyzingor sorting fluorescent and non-fluorescent nucleotide fragments, theinvention can also provide multiparameter analysis. For example, sizingor sorting can be done according to a window or threshold value, meaningthat molecules (e.g. polynucleotides) are selected based on the presenceof a signal above or below a certain value or threshold. There can alsobe several points of analysis on the same chip for multiple time coursemeasurements.

Mixing Embodiments. Besides analyzing and/or sorting molecules andparticles in a sample, the microfluidic devices of this invention arealso useful for mixing two or more different fluids. For example, thedevices of the invention can be used to mix fluids containing differentmolecules; such as different solvent molecules, molecules of a sampleand/or reagent, or for mixing molecules of a sample and a detectionprobe. In preferred embodiments, the devices are used to mix moleculessuch as nucleic acid molecules, polypeptide molecules (e.g., proteins),antibody molecules, or molecules of a particular reagent or ligand. Inother preferred embodiments the devices can be used to mix suspensionsof different particles such as cells of virions.

The invention therefore provides uses of these microfluidic devices formixing two or more different fluids and, in particular, provides methodsfor mixing different fluids using a microfluidic device that has: (i) aloop channel, (ii) at least one service channel in fluid communicationwith the loop channel, (iii) a microvalve separating the loop channelfrom the service channel, and (iv) a pump associated with the loopchannel. The methods involve introducing the different fluids to themicrofluidic device so that each different fluid is loaded into the loopchannel, and activating the pump associated with the loop channel sothat the different fluids are effectively mixed.

As used to describe the present invention, two or more fluids in amicrofluidic device or channel are said to be “effectively mixed” whenthe channel contains a homologous or substantially homologouscombination of the molecules and/or particles from the different fluids.Thus, in preferred embodiments two or more different fluids in amicrofluidic channel may be effectively mixed if the channel contains acombination of molecules and/or particles from the fluids that is, e.g.,at least 50% homologous, more preferably at least 60% homologous, 70%homologous, 75%, 80%, 85%, or 90% homologous. However, in someembodiments homologies as low as 20% or 25% will be adequate. In otherpreferred embodiments, the two or more fluids in a microfluidic deviceor channel are mixed when the combination of molecules is more than 90%homologous, more preferably 95% homologous and still more preferably atleast 99% homologous. Indeed, in particularly preferred embodiments ofthe invention, two or more fluids in a microfluidic device or channelare said to be mixed when the combination of molecules and/or particlesfrom the different fluids is 100% homologous. Thus, when two or morefluids are mixed in a microfluidic device of the invention theypreferably are no longer distinguishable as individual fluids and are,to a user, a single, homologous fluid of molecules and/or particles.

Using the microfluidic devices provided in this invention, moleculesand/or particles in fluids may be mixed in mixing times that are only afew minutes, as opposed to mixing by simple diffusion which may take amatter of hours. Thus, in preferred embodiments of the invention fluidsare mixed in a microfluidic loop by activating the pump for less thanone hour and more preferably for less than 30 minutes. In typical,preferred embodiments the pump need only be activated for 15 minutes orless (e.g., for 10, 5, 4, 3, 2 or 1 minute). In certain embodiments, themixing time may be as short as a few seconds (for example, between 60and 30 seconds, or less than 30 seconds). For example, as explained,supra, in the Examples, two fluids may be effectively mixed in amicrofluidic loop if the center of fluid front (i.e., the boundarybetween the different fluids) makes only half of one revolution throughthe loop. Typically, however, the fluid front will make multiplerevolutions around the loop (e.g., at least 1, at least 2, at least 5,at least 10, at least 50 or at least 100).

Using methods that are similar to the mixing methods described above,microfluidic devices of the invention may also be used to facilitatebinding of a sample to target. Accordingly, the invention provides suchuses for the microfluidic devices described here and, in particular,provides methods for binding a sample to a target using a microfluidicdevice which has: (i) a loop channel, (ii) at least one service channelin fluid communication with the loop channel, (iii) a microvalveseparating the loop channel from the service channel, and (iv) a pumpassociated with the loop channel. In such methods, molecules of thetarget are preferably disposed within the loop channel of themicrofluidic device. The methods therefore simply comprise introducing afluid containing the sample to the microfluidic device, and activatingthe pump so that the fluid throughs through the loop channel. Theparticles or molecules in the sample therefore travel and bind morerapidly to the target molecules than by simple diffusion so that, ingeneral, the pump need only be activated for a matter of minutes (e.g.,fewer than 60 minutes, more preferably fewer than 30, 20, 15, 10, 5, 4,3, 2 or 1 minutes). In fact, in some instances the pump may even beactivated for less than one minute (e.g., for between 60-30 seconds, orfor less than 30 seconds). The sample may be, e.g., a suspension ofparticles such as cells or virions, or that sample may be a solution ofmolecules. For example, in a preferred embodiment the sample is a sampleof nucleic acid molecules that includes (or is suspected to include) anucleotide sequence of interest to a user. In such a preferredembodiment, the target molecules are typically polynucleotide probes(e.g., having a sequence complementary to the nucleotide sequence ofinterest). In still other embodiments, the sample may comprisepolypeptide molecules and the target may comprise antibody probemolecules or, alternatively, the sample may comprise a particular ligandor reagent, and the target molecules may comprise a polypeptide orprotein probe (e.g., that binds to and/or reacts with the sample).

Thus, the invention provides for the rapid and accurate determination ofthe “profile” of a polynucleotide in high resolution using minimalamounts of material in these simple and inexpensive microfabricateddevices. The methods and devices of the invention can replace or be usedin combination with conventional gel based approaches.

The devices and methods of the invention can also be used to test asample against multiple targets. In these embodiments an array ofprobes, corresponding to a different targets, is fixed to one or morechannels in one or more detection regions of the device. Preferably,probes are fixed to a glass substrate or coverslip that seals thedetection channels while exposing them to optical interrogation orexamination, e.g. with a microscope. Sample is passed by the probes, andmatching molecules in the sample (if any) respond, e.g. by reacting,associating or binding with the corresponding probes in a detectable ormeasurable way. In a preferred embodiment the sample contains DNA (e.g.a blood sample), which may be denatured or fragmented, and the probescontain DNA that is characteristic of particular disease conditions, orhas some other characteristic of interest, e.g. forensic. Any suitableset of reagents or probes can be used to provide a battery or array oftests upon a sample. Thus, the devices and methods of the invention alsoprovide high throughput screening of samples for any purpose, includingdiagnostics and drug discovery.

In one embodiment, probes are fixed in discrete locations on a glasssubstrate in a pattern corresponding to the path of an adjacenttreatment channel, or loop, which can comprise any closed path, i.e. itcan be temporarily isolated from the rest of the chip, for example byclosing valves in any channels which lead into or out of the loop. Usingthe microvalves and peristaltic pump action described herein, sample canbe introduced to the loop, containing probes, and can be recirculatedpast the probes as desired, to rapidly and repeatedly test for thepresence or absence of multiple targets in the sample.

A typical target loop of the invention, e.g. for DNA assays, has acircular path, although any path which can be closed is encompassed bythe invention. The length of the loop (or the diameter of a circularembodiment) depends on the number of diagnostic spots (e.g.hybridization spots) in the loop, the size of each spot, and thedistance between each spot. For example, a loop channel that is 100 μmwide can be provided with diagnostic spots that are about 100 μm wideand 100 μm long (or about 100 μm in diameter), with about 100 μm betweeneach spot. Each spot can be provided with a probe, for example a DNAfragment or an antibody immobilized on a substrate and presented tosample that is circulated in the loop. The spots can be observed orimaged as described herein, to detect or measure the interaction betweenmaterial in the sample and material at the diagnostic spot.

In another preferred embodiment a target loop may have some othergeometry, for example the geometry illustrated in FIG. 20. In aparticularly preferred embodiment a microfluidic device comprises anarray of target loops having a size and dimensions comparable with thewells of a standard microtiter plate. The target loops may then beassembled over the separate wells of the microtiter plate. For example,the invention provides microfluidic devices having an array of 96 targetloops (e.g., that is compatible with a 96-well microtiter plate). Theinvention also provides microfluidic devices having arrays of 384 targetloops (e.g., that are compatible with 384-well microtiter plates). Theinvention still further provides microfluidic devices having arrays of1536 target loops (e.g., that are compatible with 1.536-well microtiterplates).

In preferred embodiments the loop channel is about 2-20 μm deep,preferably about 10 μm deep, and is from about 10-200 μm wide,preferably from about 50-100 μm wide, and more preferably about 100 μmwide. The target loop is fed by a loop inlet and is drained by a loopoutlet, each of which can be independently opened and closed, e.g. byappropriately positioned microvalves. The target loop or targettreatment channel is intersected by at least three air or controlchannels on a facing layer of the device. Preferably, the controlchannels intersect the circular loop in a radial fashion, and mayterminate inside a region defined by the loop. Each intersection betweenthe loop and a control channel forms a microvalve. Varying the pressure(e.g. air pressure) in at least three intersecting control channelscreates three microvalves which open and close in response to pressurechanges, causing a peristaltic flow around the loop. For example,expansion of the control channel in response to pressure can pinch,constricting or block the loop channel at the intersection point.Relaxation of the control channel in response to a pressure drop opensthe restricted or closed loop channel. Cycles of contraction andexpansion cause temporary closing and opening of the loop channel, whichsets up a flow around the loop, which is preferably closed duringtesting for matching probes. In this system, sample molecules arecirculated past the probes in a closed loop, ensuring rapid and completeexposure of the sample to a plurality of probes, for quick, accurate,and inexpensive analysis using very small amounts of sample, probes,reagents, etc.

4. BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows a nucleic acid sorting device in accordance with oneembodiment of the invention.

FIG. 2 shows a partial perspective view of a nucleic acid sortingdevice, showing a sample solution reservoir and sample inlet.

FIG. 3A shows one embodiment of a detection region used in a nucleicacid sorting device, having an integrated photodiode detector.

FIG. 3B shows another embodiment of a detection region, having anintegrated photodiode detector, and providing a larger detection volume(than the embodiment of FIG. 3A).

FIGS. 4A-4B show one embodiment of a valve within a branch channel of anucleic acid sorting device, and steps in fabrication of the valve.

FIG. 5A shows one embodiment of a discrimination region used in anucleic acid sorting device, having electrodes disposed within thechannels for electrophoretic discrimination.

FIG. 5B shows another embodiment of a discrimination region used in anucleic acid sorting device, having electrodes disposed forelectroosmotic discrimination.

FIGS. 5C and 5D show two further embodiments of a discrimination region,having valves disposed for pressure electrophoretic separation, wherethe valves are within the branch point, as shown in 4C, or within thebranch channels, as shown in 4D.

FIG. 6 shows a device with analysis units containing a cascade ofdetection and discrimination regions suitable for successive rounds ofpolynucleotide or cell sorting.

FIGS. 7A-7D show initial steps in photolithographic microfabrication ofa nucleic, acid sorting device from a silicon wafer, usingphotolithography and several stages of etching.

FIG. 8 shows a schematic representation of a process for obtaining asilicone elastomer impression of a silicon mold to provide amicrofabricated chip according to the invention.

FIG. 9 shows a schematic representation of an apparatus of theinvention, in which a silicone elastomer chip is mounted on an invertedmicroscope for optical detection of a laser-stimulated reporter.Electrodes are used to direct cells in response to the microscopedetection.

FIG. 10 is a photograph of an apparatus of the invention, showing a chipwith an inlet channel and reservoir, a detection region, a branch point,and two outlet channels with reservoirs.

FIGS. 11A and B show a sorting scheme according to the invention, indiagrammatic form.

FIGS. 12A and B show a reversible sorting scheme according to theinvention.

FIGS. 13A and 13B are micrographs of an exemplary chip according to theinvention, e.g. for DNA diagnosis. FIG. 13A shows input mixingT-channels 5 (50 μm wide×10 μm deep) on a lower layer and sixcorresponding air channels 7 and control microvalves 17 on an upperlayer. A wider (100-μm) air channel 3 is used to close the inlet 15 at avalve 13 when the peristaltic pump at the ring (FIG. 13B) startsoperating, e.g. for mixing or hybridization. FIG. 13B shows a centerring loop 28 for mixing and/or DNA hybridization. Any three of thefinger channels 22 (on the top air-channel layer) form a peristalticpump with corresponding valves 32. The loop channels are 50 μm wide×10μm deep. The channel 15 at the bottom of FIG. 13A connects to thechannel 15 top of FIG. 13B. The whole device is 1″ by 1″ in size.

FIG. 14 is a schematic diagram of a device of the invention, showing acentral mixing and/or a detection loop (e.g. for hybridization) actuatedby a peristaltic pump formed of microvalves where air channels intersectunderlying fluid channels.

FIG. 15 is a schematic depiction of a peristaltic pump, formed by threeair channels intersecting an underlying fluid channels, with amicrovalve at each intersection.

FIGS. 16A-C show images of chemically patterned cover slips, e.g. foruse in immobilizing probes within a detection loop of the invention. InFIG. 16A, a line pattern is obtained by flowing avidin-fluoresceinconjugates vertically on a biotinylated cover slip. In FIG. 16B acheckerboard pattern obtained by flowing strepta-vidin horizontally (200μm) and biotin-fluorescein conjugates vertically (100 μm). FIG. 16Cshows DNA patterned on a silanized slide. The DNA lights up when thefluorescent dye PicoGreen is flowing in the central ring loop. Fromtop-left to bottom-right in the figure shows DNA with a slightauto-fluorescence. From top-right to bottom-left shows part of thecentral ring of the diagnosis chip under a dark-field illumination.

FIG. 17 is a schematic diagram of a device of FIGS. 13A and 13B, showinga T-inlet and a circular mixing and detection loop with cooperatingcontrol channels forming microvalves and a peristaltic pump.

FIG. 18 shows in-line mixing by rotary pumping in a closed loop channelof the invention. In FIG. 18A there is no pumping. Buffer containingfluorescent beads (left) and buffer containing fluorescent dye (right)do not mix with each other because of a laminar flow profile. FIG. 18Bshows active rotary pumping. The peristaltic pump at the ring, where theliquids come from, is turned on. Both dye and beads are well mixed atthe output channel; Each inset shows an illustration of thecorresponding distribution of beads and fluorescent dye in the fluidicchannels.

FIG. 19 is a snapshot of a running biotin/avidin diagnosis chip, wheretwo biotin pads at the central ring are shown. This picture was taken 4minutes after the peristaltic pump started. Fluorescent NeutrAvidinbeads have bound to the biotin pads. Thin stripes are beads which werestill moving in the ring loop while this picture was taken.

FIG. 20 is a schematic diagram of an exemplary loop channel (e.g., forhybridization) geometry which may be used in a microfluidic device.

FIGS. 21A-B are schematic diagrams of a mixing or detection loop (e.g.,for hybridization) actuated by peristaltic pumps which can be used in amicrofluidic device having an array of such loops.

FIGS. 22A-B are schematic diagrams of an exemplary microfluidic devicehaving an arrayed plurality of the mixing or detection loops illustratedin FIGS. 21A-B.

FIGS. 23A-F shows images (FIGS. 23A-B) and data (FIGS. 23C-F) fromcontinuous flow mixing by rotary pumping in an open loop channel of theinvention. In FIG. 23A there is no pumping through the loop. FIGS. 23Cand 23E show the distribution of the fluorescent dye solution and beadsin detail when there is no rotary pumping in the loop. Buffer containingfluorescent beads (left) and fluorescent dye (right) do not mix witheach other because of a laminar flow profile. FIG. 23B showscontinuous-flow mixing in the loop when there is active rotary pumping.FIGS. 23D and 23F show the distribution of the fluorescent dye solutionand beads in detail when there is active rotary pumping in the loop.Rotary mixing helped about 25% of the fluorescent beads traverse intothe other side of the flow stream.

FIG. 24 is a simplified schematic diagram for an exemplary loop channelthat has a channel width of 2r₀ (i.e., a radius equal to r₀) and forms acircular loop of radius R. The loop channel is used, in this example, tomix to separate fluid indicated as fluid A and B, respectively.

5. DETAILED DESCRIPTION OF THE INVENTION 5.1. Definitions

The terms used in this specification generally have their ordinarymeanings in the art, within the context of the invention, and in thespecific context where each term is used. For example, conventionaltechniques of molecular biology, microbiology and recombinant DNAtechniques may be employed in accordance with the present invention.Such techniques and the meanings of terms associated therewith areexplained fully in the literature. See, for example, Sambrook, Fitsch &Maniatis, Molecular Cloning: A Laboratory Manual, Second Edition (1989)Cold Spring Harbor Laboratory Press, Cold Spring Harbor, N.Y. (referredto herein as “Sambrook et al., 1989”); DNA Cloning: A PracticalApproach, Volumes I and II (D. N. Glover ed. 1985); OligonucleotideSynthesis (M. J. Gait ed. 1984); Nucleic Acid Hybridization (B. D. Hames& S. J. Higgins, eds. 1984); Animal Cell Culture (R. I. Freshney, ed.1986); Immobilized Cells and Enzymes (IRL Press, 1986); B. E. Perbal, APractical Guide to Molecular Cloning (1984); F. M. Ausubel et al.(eds.), Current Protocols in Molecular Biology, John Wiley & Sons, Inc.(1994). See also, PCR Protocols: A Guide to Methods and Applications,Innis et al., eds., Academic Press, Inc., New York (1990); Saiki et al.,Science 1988, 239:487; and PCR Technology: Principles and Applicationsfor DNA Amplification, H. Erlich, Ed., Stockton Press.

Certain terms that are used to describe the invention are discussedbelow, or elsewhere in the specification, to provide additional guidanceto the practitioner in describing the devices and methods of theinvention and how to make and use them. For convenience, certain termsare highlighted, for example using italics and/or quotation marks. Theuse of highlighting has no influence on the scope and meaning of a term;the scope and meaning of a term is the same, in the same context,whether or not it is highlighted. It will be appreciated that the samething can be said in more than one way. Consequently, alternativelanguage and synonyms may be used for any one or more of the termsdiscussed herein, nor is any special significance to be placed uponwhether or not a term is elaborated or discussed herein. Synonyms forcertain terms are provided. A recital of one or more synonyms does notexclude the use of other synonyms. The use of examples anywhere in thisspecification, including examples of any terms discussed herein, isillustrative only, and in no way limits the scope and meaning of theinvention or of any exemplified term. Likewise, the invention is notlimited to the preferred embodiments.

As used herein, “about” or “approximately” shall generally mean within20 percent, preferably within 10 percent, and more preferably within 5percent of a given value or range. Numerical quantities given herein areapproximate, meaning that the term “about” or “approximately” can beinferred if not expressly stated.

The term “molecule” means any distinct or distinguishable structuralunit of matter comprising one or more atoms, and includes for examplepolypeptides and polynucleotides.

The term “polymer” means any substance or compound that is composed oftwo or more building blocks (‘mers’) that are repetitively linked toeach other. For example, a “dimer” is a compound in which two buildingblocks have been joined together.

The term “polynucleotide” as used herein refers to a polymeric moleculehaving a backbone that supports bases capable of hydrogen bonding totypical polynucleotides, where the polymer backbone presents the basesin a manner to permit such hydrogen bonding in a sequence specificfashion between the polymeric molecule and a typical polynucleotide(e.g., single-stranded DNA). Such bases are typically inosine,adenosine, guanosine, cytosine, uracil and thymidine. Polymericmolecules include double and single stranded RNA and DNA, and backbonemodifications thereof, for example, methylphosphonate linkages.

Thus, a “polynucleotide” or “nucleotide sequence” is a series ofnucleotide bases (also called “nucleotides”) generally in DNA and RNA,and means any chain of two or more nucleotides. A nucleotide sequencetypically carries genetic information, including the information used bycellular machinery to make proteins and enzymes. These terms includedouble or single stranded genomic and cDNA, RNA, any synthetic andgenetically manipulated polynucleotide, and both sense and anti-sensepolynucleotide (although only sense stands are being representedherein). This includes single- and double-stranded molecules, i.e.,DNA-DNA, DNA-RNA and RNA-RNA hybrids, as well as “protein nucleic acids”(PNA) formed by conjugating bases to an amino acid backbone This alsoincludes nucleic acids containing modified bases, for examplethio-uracil, thio-guanine and fluoro-uracil.

The polynucleotides herein may be flanked by natural regulatorysequences, or may be associated with heterologous sequences, includingpromoters, enhancers, response elements, signal sequences,polyadenylation sequences, introns, 5′- and 3′-non-coding regions, andthe like. The nucleic acids may also be modified by many means known inthe art. Non-limiting examples of such modifications includemethylation, “caps”, substitution of one or more of the naturallyoccurring nucleotides with an analog, and internucleotide modificationssuch as, for example, those with uncharged linkages (e.g., methylphosphonates, phosphotriesters, phosphoroamidates, carbamates, etc.) andwith charged linkages (e.g., phosphorothioates, phosphorodithioates,etc.). Polynucleotides may contain one or more additional covalentlylinked moieties, such as, for example, proteins (e.g., nucleases,toxins, antibodies, signal peptides, poly-L-lysine, etc.), intercalators(e.g., acridine, psoralen, etc.), chelators (e.g., metals, radioactivemetals, iron, oxidative metals, etc.), and alkylators. Thepolynucleotides may be derivatized by formation of a methyl or ethylphosphotriester or an alkyl phosphoramidate linkage. Furthermore, thepolynucleotides herein may also be modified with a label capable ofproviding a detectable signal, either directly or indirectly. Exemplarylabels include radioisotopes, fluorescent molecules, biotin, and thelike.

“DNA” (deoxyribonucleic acid) means any chain or sequence of thechemical building blocks adenine (A), guanine (G), cytosine (C) andthymine (T), called nucleotide bases, that are linked together on adeoxyribose sugar backbone. DNA can have one strand of nucleotide bases,or two complimentary strands which may form a double helix structure.“RNA” (ribonucleic acid) means any chain or sequence of the chemicalbuilding blocks adenine (A), guanine (G), cytosine (C) and uracil (U),called nucleotide bases, that are linked together on a ribose sugarbackbone. RNA typically has one strand of nucleotide bases.

A “polypeptide” (one or more peptides) is a chain of chemical buildingblocks called amino acids that are linked together by chemical bondscalled peptide bonds. A protein or polypeptide, including an enzyme, maybe “native” or “wild-type”, meaning that it occurs in nature; or it maybe a “mutant”, “variant” or “modified”, meaning that it has been made,altered, derived, or is in some way different or changed from a nativeprotein, or from another mutant.

As used herein, “cell” means any cell or cells, as well as viruses orany other particles having a microscopic size, e.g. a size that issimilar to that of a biological cell, and includes any prokaryotic oreukaryotic cell, e.g., bacteria, fungi, plant and animal cells. Cellsare typically spherical, but can also be elongated, flattened,deformable and asymmetrical, i.e., non-spherical. The size or diameterof a cell typically ranges from about 0.1 to 120 microns, and typicallyis from about 1 to 50 microns. A cell may be living or dead. Since, themicrofabricated device of the invention is directed to analyzing orsorting materials having a size similar to protein or polynucleotidemolecules or to biological cells (e.g. about 0.1 to 120 microns), anymaterial having a size similar to these molecules and cells can becharacterized and sorted using the microfabricated device of theinvention. Channels and devices of appropriate size can be fabricatedfor larger or smaller materials, e.g. for any materials of microscopicsize. Thus, the term cell shall further include microscopic beads (suchas chromatographic and fluorescent beads), liposomes, emulsions, or anyother encapsulating biomaterials and porous materials. Non-limitingexamples include latex, glass, or paramagnetic beads; and vesicles suchas emulsions and liposomes, and other porous materials such as silicabeads. Beads ranging in size from 0.1 micron to 1 mm can also be used,for example in sorting a library of compounds produced by combinatorialchemistry. As used herein, a cell may be charged or uncharged. Forexample, charged beads may be used to facilitate flow or detection, oras a reporter. Biological cells, living or dead, may be charged forexample by using a surfactant, such as SDS (sodium dodecyl sulfate).

A “reporter” is any molecule, or a portion thereof, that is detectable,or measurable, for example, by optical detection: In addition, thereporter associates with a molecule or cell or with a particular markeror characteristic of the molecule or cell, or is itself detectable, topermit identification of the molecule or cell, or the presence orabsence of a characteristic of the molecule or cell. In the case ofmolecules such as polynucleotides such characteristics include size,molecular weight, the presence or absence of particular constituents ormoieties (such as particular nucleotide sequences or restrictionssites). The term “label” can be used interchangeably with “reporter”.The reporter is typically a dye, fluorescent, ultraviolet, orchemiluminescent agent, chromophore, or radio-label, any of which may bedetected with or without some kind of stimulatory event, e.g., fluorescewith or without a reagent. Typical reporters for molecularfingerprinting include without limitation fluorescently-labeled singlenucleotides such as fluorescein-dNTP, rhodamine-dNTP, Cy3-dNTP,Cy5-dNTP, where dNTP represents dATP, dTTP, dUTP or dCTP. The reportercan also be chemically-modified single nucleotides, such as biotin-dNTP.Alternatively, chemicals can be used that react with an attachedfunctional group such as biotin.

A “marker” is a characteristic of a molecule or cell that is detectableor is made detectable by a reporter, or which may be coexpressed with areporter. For molecules, a marker can be particular constituents ormoieties, such as restrictions sites or particular nucleic acidsequences in the case of polynucleotides. The marker may be directly orindirectly associated with the reporter or can itself be a reporter.Thus, a marker is generally a distinguishing feature of a molecule, anda reporter is generally an agent which directly or indirectly identifiesor permits measurement of a marker. These terms may, however, be usedinterchangeably.

The term “flow” means any movement of liquid or solid through a deviceor in a method of the invention, and encompasses without limitation anyfluid stream, and any material moving with, within or against thestream, whether or not the material is carried by the stream. Forexample, the movement of molecules or cells through a device or in amethod of the invention, e.g. through channels of a microfluidic chip ofthe invention, comprises a flow. This is so, according to the invention,whether or not the molecules or cells are carried by a stream of fluidalso comprising a flow, or whether the molecules or cells are caused tomove by some other direct or indirect force or motivation, and whetheror not the nature of any motivating force is known or understood. Theapplication of any force may be used to provide a flow, includingwithout limitation, pressure, capillary action, electroosmosis,electrophoresis, dielectrophoresis, optical tweezers, and combinationsthereof, without regard for any particular theory or mechanism ofaction, so long as molecules or cells are directed for detection,measurement or sorting according to the invention.

An “inlet region” is an area of a microfabricated chip that receivesmolecules or cells for detection measurement or sorting. The inletregion may contain an inlet channel, a well or reservoir, an opening,and other features which facilitate the entry of molecules or cells intothe device. A chip may contain more than one inlet region if desired.The inlet region is in fluid communication with the main channel and isupstream therefrom.

An “outlet region” is an area of a microfabricated chip that collects ordispenses molecules or cells after detection, measurement or sorting. Anoutlet region is downstream from a discrimination region, and maycontain branch channels or outlet channels. A chip may contain more thanone outlet region if desired.

An “analysis unit” is a microfabricated substrate, e.g., amicrofabricated chip, having at least one inlet region, at least onemain channel, at least one detection region and at least one outletregion. Sorting embodiments of the analysis unit include adiscrimination region and/or a branch point, e.g. downstream of thedetection region, that forms at least two branch channels and two outletregions. A device of the invention may comprise a plurality of analysisunits.

A “main channel” is a channel of the chip of the invention which permitsthe flow of molecules or cells past a detection region for detection(identification), measurement, or sorting. In a chip designed forsorting, the main channel also comprises a discrimination region. Thedetection and discrimination regions can be placed or fabricated intothe main channel. The main channel is typically in fluid communicationwith an inlet channel or inlet region, which permit the flow ofmolecules or cells into the main channel. The main channel is alsotypically in fluid communication with an outlet region and optionallywith branch channels, each of which may have an outlet channel or wastechannel. These channels permit the flow of molecules or cells out of themain channel.

In certain embodiments, a “circulation loop” is located within the chip,typically in or communicating with the main channel, in which a fluid(e.g. the flow of a biological sample) is circulated. The circulationloop may comprise a “hybridization loop” or “target loop” in which theflow is directed past a series of targets or probes (e.g. DNA orproteins) that are in or exposed to the loop and its contents. Forexample, probes may be patterned on the surface of a substrate, e.g. asolid substrate and also called a “probe substrate”. The probe substratetypically forms part of a channel, e.g. as a wall, ceiling or floor of afluid channel, or is exposed to or communicates with a channel, orreceives or is exposed to a flow of fluid or sample from a channel.Preferred probe substrates are transparent, e.g. glass. Alternatively,the probe substrate may be an elastomer itself, which may proveadvantageous when higher back pressures are used.

The loop may have any shape. The channel or channels comprising a loopmay have or cooperate with pumps and/or valves to open and close theloop, and/or to provide or drain contents to and from the loop. In apreferred embodiment, the loop can be isolated or closed from otherchannels in a microfluidic device. Also in a preferred embodiment, fluidcan be circulated in the loop, for example by providing a peristalticpump comprising three or more microvalves.

A circulation loop may also be referred to as a “detection loop” inembodiments where detection, measurement or analysis occurs in orcoincident with all or any part of the loop.

A “detection region” is a location within the chip, typically in orcoincident with the main channel (or a portion thereof) and/or in orcoincident with a detection loop, where molecules or cells to beidentified, characterized, hybridized, measured, analyzed or sorted(etc.), are examined on the basis of a predetermined characteristic. Ina preferred embodiment, molecules or cells are examined one at a time.In other preferred embodiments, molecules, cells or samples are examinedtogether, for example in groups, in arrays, in rapid, simultaneous orcontemporaneous serial or parallel arrangements, or by affinitychromatography. In one such embodiment, a sample is exposed to probes indetection region, preferably probes having a predetermined patternwithin or coincident with a detection region, e.g. a targethybridization or detection loop. Preferably, the molecule or cellcharacteristic is detected or measured optically, for example, bytesting for the presence or amount of a reporter. For example, thedetection region is in communication with one or more microscopes,diodes, light stimulating devices, (e.g., lasers), photomultipliertubes, and processors (e.g., computers and software), and combinationsthereof, which cooperate to detect a signal representative of acharacteristic, marker, or reporter, and to determine and direct themeasurement or the sorting action at the discrimination region. Insorting embodiments, the detection region is in fluid communication witha discrimination region and is at, proximate to, or upstream of thediscrimination region.

A “discrimination region” or “branch point” is a junction of a channelwhere the flow of molecules or cells can change direction to enter oneor more other channels, e.g., a branch channel, depending on a signalreceived in connection with an examination in the detection region.Typically, a discrimination region is monitored and/or under the controlof a detection region, and therefore a discrimination region may“correspond” to such detection region. The discrimination region is incommunication with and is influenced by one or more sorting techniquesor flow control systems, e.g., electric, electro-osmotic, (micro-)valve, etc. A flow control system can employ a variety of sortingtechniques to change or direct the flow of molecules or cells into apredetermined branch channel.

A “branch channel” is a channel which is in communication with adiscrimination region and a main channel. Typically, a branch channelreceives molecules or cells depending on the molecule or cellcharacteristic of interest as detected by the detection region andsorted at the discrimination region. A branch channel may be incommunication with other channels to permit additional sorting.Alternatively, a branch channel may also have an outlet region and/orterminate with a well or reservoir to allow collection or disposal ofthe molecules or cells.

The term “forward sorting” or flow describes a one-direction flow ofmolecules or cells, typically from an inlet region (upstream) to anoutlet region (downstream), and preferably without a change indirection, e.g., opposing the “forward” flow. Preferably, molecules orcells travel forward in a linear fashion, i.e., in single file. Apreferred “forward” sorting algorithm consists of running molecules orcells from the input channel to the waste channel, until a molecule orcell is identified to have an optically detectable signal (e.g.fluorescence) that is above a pre-set threshold, at which point voltagesare temporarily changed to electroosmotically divert the molecule or tothe collection channel.

The term “reversible sorting” or flow describes a movement or flow thatcan change, i.e., reverse direction, for example, from a forwarddirection to an opposing backwards direction. Stated another way,reversible sorting permits a change in the direction of flow from adownstream to an upstream direction. This may be useful for moreaccurate sorting, for example, by allowing for confirmation of a sortingdecision, selection of particular branch channel, or to correct animproperly selected channel.

Different “sorting algorithms” can be implemented in devices of theinvention by different programs or protocols, for example under thecontrol of a personal computer. A “sorting algorithm” is any set ofsteps by which any items are identified, distinguished or separated. Asan example, consider a pressure-switched scheme instead ofelectro-osmotic flow. Electro-osmotic switching is virtuallyinstantaneous and throughput is limited by the highest voltage that canbe applied to the sorter (which also affects the run time through iondepletion effects). A pressure switched-scheme does not require highvoltages and is more robust for longer runs. However, mechanicalcompliance in the system is likely to cause the fluid switching speed tobecome rate-limiting with the “forward” sorting program. Since the fluidis at low Reynolds number and is completely reversible, when trying toseparate rare molecules or cells one can implement a sorting algorithmthat is not limited by the intrinsic switching speed of the device. Themolecules or cells flow at the highest possible static (non-switching)speed from the input to the waste. When an interesting molecule or cellis detected, the flow is stopped. By the time the flow stops, themolecule or cell may be past the junction and part way down the wastechannel. The system is then run backwards at a slow (switchable) speedfrom waste to input, and the molecule or cell is switched to thecollection channel when it passes through the detection region. At thatpoint, the molecule or cell is “saved” and the device can be run at highspeed in the forward direction again. Similarly, an device of theinvention that is used for analysis, without sorting, can be run inreverse to re-read or verify the detection or analysis made for one ormore molecules or cells in the detection region. This “reversible”analysis or sorting method is not possible with standard gelelectrophoresis technologies (for molecules) nor with conventional FACSmachines (for cells). Reversible algorithms are particularly useful forcollecting rare molecules or cells or making multiple time coursemeasurements of a molecule or single cell.

A “gene” is a sequence of nucleotides which code for a functionalpolypeptide. For the purposes of the invention a gene includes an mRNAsequence which may be found in the cell. For example, measuring geneexpression levels according to the invention may correspond to measuringmRNA levels. “Genomic sequences” are the total set of genes in aorganism. The term “genome” denotes the coding sequences of the totalgenome.

Polynucleotides are “hybridizable” to each other when at least onestrand of one polynucleotide can anneal to another polynucleotide underdesired or defined stringency conditions. Stringency of hybridization isdetermined, e.g., by a) the temperature at which hybridization and/orwashing is performed, and b) the ionic strength and polarity (e.g.,formamide) of the hybridization and washing solutions, as well as otherparameters. Hybridization requires that the two polynucleotides containsubstantially complementary sequences; depending on the stringency ofhybridization, however, mismatches may be tolerated. Typically,hybridization of two sequences at high stringency (such as, for example,in an aqueous solution of 0.5×SSC at 65° C.) requires that the sequencesexhibit some high degree of complementarity over their entire sequence.Conditions of intermediate stringency (such as, for example, an aqueoussolution of 2×SSC at 65° C.) and low stringency (such as, for example,an aqueous solution of 2×SSC at 55° C.), require correspondingly lessoverall complementarity between the hybridizing sequences. (1×SSC is0.15 M NaCl, 0.015 M Na citrate.) Polynucleotides that “hybridize” tothe polynucleotides herein may be of any length. In one embodiment, suchpolynucleotides are at least 10, preferably at least 15 and mostpreferably at least 20 nucleotides long. In another embodiment,polynucleotides that hybridizes are of about the same length. In anotherembodiment, polynucleotides that hybridize include those which annealunder suitable stringency conditions and which encode polypeptides orenzymes having the same function.

5.2. Overview of the Invention

The invention provides devices and methods for the detection of multiplediseases in humans or animals. More particularly, in the microfabricateddevice according to the invention, detection of the presence ofmolecules (i.e., polynucleotides, proteins, or antigen/antibodycomplexes) are correlated to a hybridization signal from anoptically-detectable (e.g. fluorescent) reporter associated with thebound molecules. The polynucleotides may also be fragmented, for exampleusing endonucleases, to produce a set of fragments that vary in size.The size distribution of these fragments (e.g. the number of fragmentsof each size over a range of sizes) may uniquely identify the source ofthe sample. Some or all of the fragments can be selected to serve as a“fingerprint” of the sample. Further, fragments comprising thefingerprint can be labeled, for example with a reporter molecule such asfluorescent marker, so that the they can be more readily detected,measured or sorted. Universal chips according to the invention can befabricated not only with DNA but also with other molecules such as RNA,proteins, peptide nucleic acid (PNA) and polyamide molecules (4), toname a few.

Thus, the invention provides rapid and accurate determination of thepresence of particular genes correlated to a particular disease usingminimal amounts of a sample in simple and inexpensive microfabricateddevices. The methods and devices of the invention can replace or be usedin combination with conventional gel based approaches.

These measurements can be detected by any suitable means, preferablyoptical, and can be stored for example in a computer as a representationof the presence or absence of a particular gene or the fragmentscomprising the fingerprint of that gene. Depending on the strategy forproducing fragments which comprise a fingerprint, oligonucleotide probesof known composition and length may be used to “tag” or label thefragments. For example, probes having sequences that are complementaryto each of the fragments can be made by combining the fragments withlabeled nucleotide bases in the presence of a polymerase, which is anenzyme that assembles a single strand of complementary polynucleotideusing another strand (i.e. a fingerprint fragment) as a template. Thenucleotide bases used to make these probes may be radioactive, or can belabeled with a fluorescent marker, or with some other readily detectablereporter. The resulting probes can be used to record a fingerprint ofthe sample, by detecting and measuring the lever of reporter as anindication of size, or by sorting the probes according to size.

Labeled or unlabeled probes can also be used to “fish out” matchingpolynucleotides from a test sample containing unknown DNA orpolynucleotides. Under appropriate hybridizing conditions, probes willbind to matching fragments in a sample. This can provide a way to testfor a match, for example when the probes comprising a fingerprinthybridize to complementary fragments in the sample. In a preferredembodiment, probes are immobilized on a substrate that forms part of oris exposed to fluid or treatment channels in a detection region of amicrofluidic device, e.g. a target loop having discrete hybridizationspots. The loop can be selectively isolated from the microfluidicdevice, for fluid circulation to expose samples and probes to eachother. Circulation is preferably provided by microvavles forming aperistaltic pump.

In one aspect of the invention, polynucleotides, e.g., DNA, can bedetected, sized or sorted dynamically in a continuous flow stream ofmicroscopic dimensions based for example on molecular weight, using amicrofabricated polynucleotide sorting device. The polynucleotides,suspended in a suitable carrier fluid (e.g., ddH₂0 or TE), areintroduced into an inlet end of a narrow channel in the sorting device.The molecular weight of each molecule is calculated from the intensityof signal from an optically-detectable reporter incorporated into orassociated with the polynucleotide molecule as the molecule passesthrough a “detection window” or “detection region” in the device.

In a sorter embodiment, molecules having a molecular weight fallingwithin a selected range are diverted into a selected output or “branch”channel of the device. The sorted polynucleotide molecules may becollected from the output channels and used in subsequent manipulations.

According to another aspect of the invention, a device such as describedabove, but not necessarily including components for sorting themolecules, can be used to measure or quantify the size range ofpolynucleotides in a sample, and store or feed this information into aprocessor or computer for subsequent analysis or display, e.g., as asize distribution histogram. Such a device enables the generation of thetype of polynucleotide fragment length data now commonly obtained fromanalytical gels, such as agarose or polyacrylamide gels, or fromSouthern blot results, in a fraction of the time required forpreparation and analysis of gels, and using a substantially smalleramount of sample.

5.2.1. Microfabricated Microfluidic Chip Architecture and Method

A molecular or cell analyzer or sorter according to the inventioncomprises at least one analysis unit having an inlet region incommunication with a main channel, a target loop, e.g. for probehybridization, a detection region within or coincident with all or aportion of the main channel or target loop, and a detector associatedwith the detection region. Sorter embodiments also have a discriminationregion or branch point in communication with the main channel and withbranch channels, and a flow control responsive to a detector. There maybe a plurality of detection regions and detectors, working independentlyor together, to analyze one or more properties of a sample. The branchchannels may each lead to an outlet region and to a well or reservoir.The inlet region may also communicate with a well or reservoir. As eachmolecule or cell passes into the detection region, it is examined for apredetermined characteristic (i.e. using the detector), and acorresponding signal is produced, for example indicating that “yes” thecharacteristic is present, or “no” it is not. The signal may correspondto a characteristic qualitatively or quantitatively. That is, the amountof the signal can be measured and can correspond to the degree to whicha characteristic is present. For example, the strength of the signal mayindicate the size of a molecule, or the potency or amount of an enzymeexpressed by a cell, or a positive or negative reaction such as bindingor hybridization of one molecule with another. In response to thesignal, data can be collected and/or a flow control can be activated todivert a molecule or cell into one branch channel or another. Thus,molecules or cells within a discrimination region can be sorted into anappropriate branch channel according to a signal produced by thecorresponding examination at a detection region. Optical detection ofmolecule or cell characteristics is preferred, for example directly orby use of a reporter associated with a characteristic chosen forsorting. However, other detection techniques may also be employed.

A variety of channels for sample flow and mixing can be microfabricatedon a single chip and can be positioned at any location on the chip asdetection and discrimination or sorting points, e.g., for kineticstudies (24, 26). A plurality of analysis units of the invention may becombined in one device. Microfabrication applied according to theinvention eliminates the dead time occurring in conventional gelelectrophoresis or flow cytometric kinetic studies, and achieves abetter time-resolution. Furthermore, linear arrays of channels on asingle chip, i.e., a multiplex system, can simultaneously detect andsort a sample by using an array of photomultiplier tubes (PMT) forparallel analysis of different channels (27). This arrangement can beused to improve throughput or for successive sample enrichment, and canbe adapted to provide a very high throughput to the microfluidic devicesthat exceeds the capacity permitted by conventional flow sorters.Circulation systems, particularly rotary circulation within a closedloop, e.g. for detection, can be used in cooperation with these andother features of the invention. Microfluidic pumps and valves are apreferred way of controlling fluid and sample flow. Microfabricationpermits other technologies to be integrated or combined on a singlechip, such as PCR (10), moving molecules or cells using opticaltweezer/trapping (28-30), transformation of cells by electroporation(31), μTAS (33), and DNA hybridization (18). Detectors and/or lightfilters that are used to detect molecule or cell characteristics ortheir reporters can also be fabricated directly on the chip.

A device of the invention can be microfabricated with a sample solutionreservoir or well at the inlet region, which is typically in fluidcommunication with an inlet channel. A reservoir may facilitateintroduction of molecules or cells into the device and into the sampleinlet channel of each analysis unit. An inlet region may have anopening, such as in the floor of the microfabricated chip, to permitentry of the sample into the device. The inlet region may also contain aconnector adapted to receive a suitable piece of tubing, such as liquidchromatography or HPLC tubing, through which a sample may be supplied.Such an arrangement facilitates introducing the sample solution underpositive pressure in order to achieve a desired flow rate through thechannels. Outlet channels and wells can be similarly provided.

5.2.2. Substrate and Flow Channels

A typical analysis unit of the invention comprises an inlet region thatis part of and feeds or communicates with a main channel, which in turncommunicates with an outlet or with two (or more) branch channels at ajunction or branch point, forming for example a T-shape or a Y-shape forsorting. Other shapes and channel geometries may be used as desired. Theregion at or surrounding the junction can also be referred to as adiscrimination region, however, precise boundaries for thediscrimination region are not required. A detection region is identifiedwithin, communicating, or coincident with a portion of the main channeldownstream of the inlet region, and in sorting embodiments, at orupstream of the discrimination region or branch point. Preciseboundaries for the detection region are not required, but are preferred.The discrimination region may be located immediately downstream of thedetection region, or it may be separated by a suitable distanceconsistent with the size of the molecules, the channel dimensions, andthe detection system. It will be appreciated that the channels can haveany suitable shape or cross-section, such as tubular or grooved, and canbe arranged in any suitable manner, so long as a flow can be directedfrom inlet to outlet, and from one channel into another, e.g. into atleast one of two or more branch channels.

The channels of the invention are microfabricated, for example byetching a silicon chip using conventional photolithography techniques,or using a micromachining technology called “soft lithography”,developed in the late 1990's (23). These and other micro fabricationmethods may be used to provide inexpensive miniaturized devices, and inthe case of soft lithography, can provide robust devices havingbeneficial properties such as improved flexibility, stability, andmechanical strength. When optical detection is employed, the inventionalso provides minimal light scatter from molecule or cell suspension andchamber material. Devices according to the invention are relativelyinexpensive and easy to set up. They can also be disposable, whichgreatly relieves many of the concerns of gel electrophoresis (formolecules) and for sterilization and permanent adsorption of particlesunto the flow chambers and channels of conventional FACS machines (forcells). Using these kinds of techniques, microfabricated fluidic devicescan replace the conventional gel electrophoresis and fluidic flowchambers of the prior art.

A microfabricated device of the invention is preferably fabricated froma silicon microchip or silicon elastomer. The dimensions of the chip arethose of typical microchips, ranging between about 0.5 cm to about 5 cmper side and about 1 micron to about 1 cm in thickness. A typical deviceof the invention is one square inch in area. The device contains atleast one analysis unit having a main channel with a centralhybridization loop and a coincident detection region. Preferably thedevice also contains at least one inlet region (which may contain aninlet channel) and one or more outlet regions (which may have fluidcommunication with a branch channel in each region). In a sortingembodiment, at least one detection region cooperates with at least onediscrimination point to divert flow via a detector-originated signal. Itshall be appreciated that the “regions” and “channels” are in fluidcommunication with each other, and therefore they may overlap, i.e.,there may be no clear boundary where a region or channel begins or ends.A microfabricated device can be transparent and can be covered with amaterial having transparent properties, e.g., a glass coverslip topermit detection of a reporter for example by an optical device, such asan optical microscope.

The dimensions of the channels and in particular of the detection regionare influenced by the size of the molecules or cells under study. Forpolynucleotides, which are large by molecular standards, a typicallength or diameter is about 3.4 angstroms per base pair. Thus, a DNA 49kpbs long, such as Lambda phage DNA, is about 17 microns long when fullyextended. A typical range of sizes for polynucleotides of the inventionis from about 1 to about 200 kpbs, or about 0.3 to about 70 microns.Detection regions used for detecting molecules have a cross-sectionalarea large enough to allow a desired molecule to pass through withoutbeing substantially slowed down relative to the flow of the solutioncarrying it. At the small dimensions of preferred embodiments of theinvention, e.g. channels of about 100 μm×10 μm, the Reynolds number isless than one, meaning that there is little or no turbulence.Nevertheless, to avoid “bottlenecks” and/or turbulence, and inembodiments where it is desirable to promote single-file flow, thechannel dimensions, particularly in the detection region, shouldgenerally be at least about twice, preferably at least about five timesas large per side or in diameter as the diameter of the largest moleculethat will be passing through it.

For molecules such as DNA, the channels in a device are between about 2to about 5 microns in width and between about 2 and about 4 or 5 micronsin depth. Similarly, the volume of the detection region in a molecularanalysis or sorting device is in the range of between about 10 to about5000 femtoliters (fl), preferably about 40 or 50 fl to about 1000 or2000 fl, most preferably on the order of about 200 fl. In preferredembodiments, the channels of the device, and particularly the channelsof a target or detection loop, are preferably between about 10 μm andabout 200 μm in width, typically 50-100 μm, and most preferably about100 μm. The channels are preferably about 2-20 μm in depth for DNA orpolynucleotide analysis, more typically about 10 μm. The detectionregion in preferred embodiments, e.g. the volume of a target loop, isbetween about 1 pl and about 1 nl.

To prevent material from adhering to the sides of the channels, thechannels (and coverslip, if used) may have a coating which minimizesadhesion. Such a coating may be intrinsic to the material from which thedevice is manufactured, or it may be applied after the structuralaspects of the channels have been microfabricated. “TEFLON” is anexample of a coating that has suitable surface properties.

A silicon substrate containing the microfabricated flow channels andother components is preferably covered and sealed, most preferably witha transparent cover, e.g., thin glass or quartz, although other clear oropaque cover materials may be used. When external radiation sources ordetectors are employed, the detection region is covered with a clearcover material to allow optical access to the molecules or cells. Forexample, anodic bonding to a “PYREX” cover slip can be accomplished bywashing both components in an aqueous H₂SO₄/H₂O₂ bath, rinsing in water,and then, for example, heating to about 350 degrees C. while applying avoltage of 450V.

5.2.3. Switching and Flow Control

Electro-osmotic and pressure-driven flow are examples of methods orsystems for flow control, that is, manipulating the flow of moleculescells, particles or reagents in one or more directions and/or into oneor more channels of a microfluidic device of the invention (20, 24, 25,34). Other methods may also be used, for example, electrophoresis anddielectrophoresis. In certain embodiments of the invention, the flowmoves in one “forward” direction, e.g. from the inlet region through themain and branch channels to an outlet region. In other embodiments thedirection of flow is reversible. Application of these techniquesaccording to the invention provides more rapid and accurate devices andmethods for analysis or sorting, for example, because the sorting occursat or in a discrimination region that can be placed at or immediatelyafter a detection region. This provides a shorter distance for moleculesor cells to travel, they can move more rapidly and with less turbulence,and can more readily be moved, examined, and sorted in single file,i.e., one at a time. In a reversible embodiment, potential sortingerrors can be avoided, for example by reversing and slowing the flow tore-read or resort a molecule or cell (or a plurality thereof) beforeirretrievably committing the molecule or cell to the outlet or to aparticular branch channel.

Without being bound by any theory, electro-osmosis is believed toproduce motion in a stream containing ions, e.g. a liquid such as abuffer, by application of a voltage differential or charge gradientbetween two or more electrodes. Neutral (uncharged) molecules or cellscan be carried by the stream. Electro-osmosis is particularly suitablefor rapidly changing the course, direction or speed of flow.Electrophoresis is believed to produce movement of charged objects in afluid toward one or more electrodes of opposite charge, and away fromone on or more electrodes of like charge. Because of its charged nature(2 charges for each base pair) DNA can be conveniently moved byelectrophoresis in a buffer of appropriate pH.

Dielectrophoresis is believed to produce movement of dielectric objects,which have no net charge, but have regions that are positively ornegatively charged in relation to each other. Alternating,non-homogeneous electric fields in the presence of particles, such asmolecules, cells or beads, cause them to become electrically polarizedand thus to experience dielectrophoretic forces. Depending on thedielectric polarizability of the particles and the suspending medium,dielectric particles will move either toward the regions of high fieldstrength or low field strength. For example, the polarizability ofliving cells depends on their composition, morphology, and phenotype andis highly dependent on the frequency of the applied electrical field.Thus, cells of different types and in different physiological statesgenerally possess distinctly different dielectric properties, which mayprovide a basis for cell separation, e.g., by differentialdielectrophoretic forces. According to formulas provided in Fiedler etal. (25), individual manipulation of single particles requires fielddifferences with dimensions close to the particles.

Manipulation is also dependent on permittivity (a dielectric property)of the particles with the suspending medium. Thus, polymer particles andliving cells show negative dielectrophoresis at high-field frequenciesin water. For example, dielectrophoretic forces experienced by a latexsphere in a 0.5 MV/m field (10V for a 20 micron electrode gap) in waterare predicted to be about 0.2 piconewtons (pN) for a 3.4 micron latexsphere to 15 pN for a 15 micron latex sphere (25). These values aremostly greater than the hydrodynamic forces experienced by the sphere ina stream (about 0.3 pN for a 3.4 micron sphere and 1.5 pN for a 15micron sphere). Therefore, manipulation of individual cells or particlescan be accomplished in a streaming fluid, such as in a cell sorterdevice, using dielectrophoresis. Using conventional semiconductortechnologies, electrodes can be microfabricated onto a substrate tocontrol the force fields in a microfabricated sorting device of theinvention. Dielectrophoresis is particularly suitable for moving objectsthat are electrical conductors. The use of AC current is preferred, toprevent permanent alignment of ions. Megahertz frequencies are suitableto provide a net alignment, attractive force, and motion over relativelylong distances. E.g. Benecke (60).

Optical tweezers can also be used in the invention to trap and moveobjects, e.g. molecules or cells, with focused beams of light such aslasers. Flow can also be obtained and controlled by providing a pressuredifferential or gradient between one or more channels of a device or ina method of the invention.

Molecules or cells can be moved by direct mechanical switching, e.g.with on-off valves, or by squeezing the channels. Pressure control mayalso be used, for example by raising or lowering an output well tochange the pressure inside the channels on the chip. See e.g. thedevices and methods described in pending U.S. application Ser. No.08/932,774 filed Sep. 25, 1997; No. 60/108,894 filed Nov. 17, 1998; No.60/086,394 filed May 22, 1998; and Ser. No. 09/325,667 filed May 21,1999 (molecular analysis systems). These methods and devices can furtherbe used in combination with the methods and devices described in pendingU.S. application Ser. No. 60/141,503 filed Jun. 28, 1999; No. 60/147,199filed Aug. 3, 1999 and Ser. No. 60/186,856 filed Mar. 3, 2000 entitled“Microfabricated Elastomeric Valve and Pump Systems”. Each of thesereferences is hereby incorporated by reference in its entirety.

Different switching and flow control mechanisms can be combined on onechip or in one device and can work independently or together as desired.

5.2.4. Detection and Discrimination for Sorting

The detector can be any device or method for interrogating a molecule orcell as it passes through the detection region. Typically, molecules orcells are to be analyzed or sorted according to a predeterminedcharacteristic that is directly or indirectly detectable, and thedetector is selected or adapted to detect that characteristic. Apreferred detector is an optical detector, such as a microscope, whichmay be coupled with a computer and/or other image processing orenhancement devices to process images or information produced by themicroscope using known techniques. For example, molecules can be sortedby size or molecular weight. Cells can be sorted for whether theycontain or produce a particular protein, by using an optical detector toexamine each cell for an optical indication of the presence or amount ofthat protein. The protein may itself be detectable, for example by acharacteristic fluorescence, or it may be labeled or associated with areporter that produces a detectable signal when the desired protein ispresent, or is present in at least a threshold amount. There is no limitto the kind or number of molecule or cell characteristics that can beidentified or measured using the techniques of the invention, whichinclude without limitation surface characteristics of the cell andintracellular characteristics, provided only that the characteristic orcharacteristics of interest for sorting can be sufficiently identifiedand detected or measured to distinguish cells having the desiredcharacteristic(s) from those which do not. For example, any label orreporter as described herein can be used as the basis for sortingmolecules or cells, i.e. detecting them to be collected.

In preferred embodiments, the molecules or cells are analyzed and/orseparated based on the intensity of a signal from anoptically-detectable reporter bound to or associated with them as theypass through a detection window or “detection region” in the device.Molecules or cells having an amount or level of the reporter at aselected threshold or within a selected range can be diverted into apredetermined outlet or branch channel of the device. The reportersignal is collected by a microscope and measured by a photomultipliertube (PMT). A computer digitizes the PMT signal and controls the flowvia valve action or electro-osmotic potentials. Alternatively, thesignal can be recorded or quantified, as a measure of the reporterand/or its corresponding characteristic or marker, e.g. for purposes ofevaluation without necessarily proceeding to sort the molecules orcells.

In one embodiment, the chip is mounted on an inverted opticalmicroscope. Fluorescence produced by a reporter is excited using a laserbeam focused on molecules (e.g. DNA) or cells passing through adetection region. Fluorescent reporters include, e.g., rhodamine,fluorescein, Texas red, Cy3, Cy5, phycobiliprotein, green fluorescentprotein (GFP), YOYO-1, and PicoGreen. In molecular fingerprintingapplications, the reporter labels are preferably a fluorescently-labeledsingle nucleotides, such as fluorescein-dNTP, rhodamine-dNTP, Cy3-dNTP,Cy5-dNTP, where dNTP represents dATP, dTTP, dUTP or dCTP. The reportercan also be chemically-modified single nucleotides, such as biotin-dNTP.Thus, in one aspect of the invention, the device can determine the sizeor molecular weight of molecules such as polynucleotide fragmentspassing through the detection region, or the presence or degree of someother characteristic indicated by a reporter. If desired, the moleculescan be sorted based on this analysis.

To detect a reporter or determine whether a molecule has a desiredcharacteristic, the detection region may include an apparatus forstimulating a reporter for that characteristic to emit measurable lightenergy, e.g., a light source such as a laser, laser diode,high-intensity lamp, (e.g., mercury lamp), and the like. In embodimentswhere a lamp is used, the channels are preferably shielded from light inall regions except the detection region. In embodiments where a laser isused, the laser can be set to scan across a set of detection regionsfrom different analysis units. In addition, laser diodes may bemicrofabricated into the same chip that contains the analysis units.Alternatively, laser diodes may be incorporated into a second chip(i.e., a laser diode chip) that is placed adjacent to themicrofabricated sorter chip such that the laser light from the diodesshines on the detection region(s).

In preferred embodiments, an integrated semiconductor laser and/or anintegrated photodiode detector are included on the silicon wafer in thevicinity of the detection region. This design provides the advantages ofcompactness and a shorter optical path for exciting and/or emittedradiation, thus minimizing distortion.

5.2.5. Sorting Schemes

According to the invention, molecules or cells are sorted dynamically ina flow stream of microscopic dimensions, based on the detection ormeasurement of a characteristic, marker or reporter that is associatedwith the molecules or cells. The stream is typically but not necessarilycontinuous, and may be stopped and started, reversed, or changed inspeed. Prior to sorting, a liquid that does not contain sample moleculesor cells can be introduced at an inlet region of the chip (e.g., from aninlet well or channel) and is directed through the device by capillaryaction, to hydrate and prepare the device for sorting. If desired, thepressure can be adjusted or equalized for example by adding buffer to anoutlet region. The liquid typically is an aqueous buffer solution, suchas ultrapure water (e.g., 18 mega ohm resistivity, obtained for exampleby column chromatography), ultrapure water, 10 mM Tris HCL and 1 mM EDTA(TE), phosphate buffer saline (PBS), and acetate buffer. Any liquid orbuffer that is physiologically compatible with the population ofmolecules or cells to be sorted can be used.

A sample solution containing a mixture or population of molecules orcells in a suitable carrier fluid (such as a liquid or buffer describedabove) is supplied to the inlet region. The capillary force causes thesample to enter the device. The force and direction of flow can becontrolled by any desired method for controlling flow, for example, by apressure differential, by valve action, or by electro-osmotic flow,e.g., produced by electrodes at inlet and outlet channels. This permitsthe movement of the molecules or cells into one or more desired branchchannels or outlet regions.

A “forward” sorting algorithm, according to the invention, includesembodiments where molecules or cells from an inlet channel flow throughthe device to a predetermined branch or outlet channel (which can becalled a “waste channel”), until the level of measurable reporter isabove a pre-set threshold. At that time, the flow is diverted to deliverthe molecule or cell to another channel. For example, in anelectro-osmotic embodiment, where switching is virtually instantaneousand throughput is limited by the highest voltage, the voltages aretemporarily changed to divert the chosen molecule or cell to anotherpredetermined outlet channel (which can be called a “collectionchannel”). Sorting, including synchronizing detection of a reporter anddiversion of the flow, can be controlled by various methods includingcomputer or microprocessor control. Different algorithms for sorting inthe microfluidic device can be implemented by different computerprograms, such as programs used in conventional FACS devices for sortingcells. For example, a programmable card can be used to controlswitching, such as a Lab PC 1200 Card, available from NationalInstruments, Austin, Tex. Algorithm's as sorting procedures can beprogrammed using C++, LABVIEW, or any suitable software. The method isadvantageous, for example, because conventional gel electrophoresismethods are generally not automated or under computer control.

A “reversible” sorting algorithm can be used in place of a “forward”mode, for example in embodiments where switching speed may be limited.For example, a pressure-switched scheme can be used instead ofelectro-osmotic flow and does not require high voltages and may be morerobust for longer runs. However, mechanical constraints may cause thefluid switching speed to become rate-limiting. In a pressure-switchedscheme the flow is stopped when a molecule or cell of interest isdetected. By the time the flow stops, the molecule or cell may be pastthe branch point and be part-way down the waste channel. In thissituation, a reversible embodiment can be used. The system can be runbackwards at a slower (switchable) speed (e.g., from waste to inlet),and the molecule or cell is then switched to a different channel. Atthat point, a potentially mis-sorted molecule or cell is “saved”, andthe device can again be run at high speed in the forward direction. This“reversible” sorting method is not possible with standard FACS machinesor in gel electrophoresis technologies. FACS machines mostly sortaerosol droplets which cannot be reversed back to the chamber, in orderto be redirected. The aerosol droplet sorter are virtually irreversible.In gel electrophoresis, molecules such as polynucleotides are drawnthrough a gel by an electric current and migrate at different ratesproportional to their molecular weights. Individual molecules can not bereversed through the gel, and indeed, altering the rate or direction ofmigration would prevent meaningful use of the technique. Reversiblesorting is particularly useful for identifying rare molecules or cells(e.g., in molecular evolution and cancer cytological identification), ormolecules or cells that are few in number, which may be misdirected dueto a margin of error inherent to any fluidic device. The reversiblenature of the device of the invention permits a reduction in thispossible error.

A “reversible” sorting method permits multiple time course measurementsof a single molecule or cell. This allows for observations ormeasurements of the same molecule or cell at different times, becausethe flow reverses the molecule or cell back into the detection windowbefore directing it to a downstream channel. Measurements can becompared or confirmed, and changes in molecule or cell properties overtime can be examined, for example in kinetic studies.

When trying to separate molecules or cells in a sample at a very lowratio to the total number of molecules or cells, a sorting algorithm canbe implemented that is not limited by the intrinsic switching speed ofthe device. Consequently, the molecules or cells flow at the highestpossible static (non-switching) speed from the inlet channel to thewaste channel. Unwanted molecules or cells can be directed into thewaste channel at the highest speed possible, and when a desired moleculeor cell is detected, the flow can be slowed down and then reversed, todirect it back into the detection region, from where it can beredirected (i.e. to accomplish efficient switching). Hence the moleculesor cells can flow at the highest possible static speed.

Preferably, the fluid carrying the molecules or cells has a relativelylow Reynolds Number, for example 10⁻². The Reynolds Number represents aninverse relationship between the density and velocity of a fluid and itsviscosity in a channel of given length. More viscous, less dense, slowermoving fluids over a shorter distance will have a lower Reynolds Number,and are easier to divert, stop, start, or reverse without turbulence.Because of the small sizes and slow velocities, microfabricated fluidsystems are often in a low Reynolds number regime (<<1). In this regime,inertial effects, which cause turbulence and secondary flows, arenegligible; viscous effects dominate the dynamics. These conditions areadvantageous for sorting, and are provided by microfabricated devices ofthe invention. Accordingly the microfabricated devices of the inventionare preferably if not exclusively operated at a low or very lowReynold's number. Exemplary sorting schemes are shown diagrammaticallyin FIGS. 11A and B and FIGS. 12A and B.

6. EXAMPLES 6.1. Microfabricated Polynucleotide Sorting Device

FIG. 1 shows an embodiment of a microfabricated polynucleotide sortingdevice 20 in accordance with the invention. The device is preferablyfabricated from a silicon microchip 22. The dimensions of the chip arethose of typical microchips, ranging between about 0.5 cm to about 5 cmper side and about 0.1 mm to about 1 cm in thickness. The devicecontains a solution inlet 24, two or more solution outlets, e.g. outlets26 and 28, and at least one analysis unit, such as the unit at 30.

Each analysis unit includes a main channel 32 having at one end a sampleinlet 34, and downstream of the sample inlet, a detection region 36, anddownstream of the detection region 36 a discrimination region 38. Aplurality of branch channels, such as channels 40 and 42, are in fluidcommunication with and branch out from the discrimination region. Thedimensions of the main and branch channels are typically between about 1μm and 10 μm per side, but may vary at various points to facilitateanalysis, sorting and/or collection of molecules.

In embodiments such as shown in FIG. 1, where the device contains aplurality of analysis units, the device may further contain collectionmanifolds, such as manifolds 44 and 46, to facilitate collection ofsample from corresponding branch channels of different analysis unitsfor routing to the appropriate solution outlet. The manifolds arepreferably microfabricated into different levels of the device, asindicated by the dotted line representing manifold 46. Similarly, suchembodiments may include a sample solution reservoir, such as reservoir48, to facilitate introduction of sample into the sample inlet of eachanalysis unit.

Also included with the device is a processor, such as processor 50. Theprocessor can be integrated into the same chip as contains the analysisunit(s), or it can be separate, e.g., an independent microchip connectedto the analysis unit-containing chip via electronic leads, such as leads52 (connected to the detection region(s) and 54 (connected to thediscrimination region(s)).

As mentioned above, the device may be microfabricated with a samplesolution reservoir to facilitate introduction of a polynucleotidesolution into the device and into the sample inlet of each analysisunit. With reference to FIG. 2, the reservoir is microfabricated intothe silicon substrate of the chip 62, and is covered, along with thechannels (such as main channel 64) of the analysis units, with a glasscoverslip 66. The device solution inlet comprises an opening 68 in thefloor of the microchip. The inlet may further contain a connector 70adapted to receive a suitable piece of tubing, such as liquidchromatography or HPLC tubing, through which the sample may be supplied.Such an arrangement facilitates introducing the sample solution underpositive pressure, to achieve a desired flow rate through the channelsas described below.

Downstream of the sample inlet of the main channel of each analysis unitis the detection region, designed to detect the level of anoptically-detectable reporter associated with polynucleotides present inthe region. Exemplary embodiments of detection regions in devices of theinvention are shown in FIGS. 3A and 3B.

6.2. Photodiode Detectors

With reference to FIG. 3A, each detection region is formed of a portionof the main channel of an analysis unit and a photodiode, such asphotodiode 72, located in the floor of the main channel. In thisembodiment, the area detectable by the detection region is the circularportion each channel defined by the receptive field of the photodiode inthat channel. The volume of the detection region is the volume of acylinder with a diameter equal to the receptive field of the photodiodeand a height equal to the depth of the channel above the photodiode.

The signals from the photodiodes are carried via output lines 76 to theprocessor (not shown), which processes the signals into valuescorresponding to the length of the polynucleotide giving rise to thesignal. The processor then uses this information, for example, tocontrol active elements in the discrimination region. The processor mayprocess the signals into values for comparison with a predetermined orreference set of values for analysis or sorting.

When more than one detection region is used, the photodiodes in thelaser diode chip are preferably spaced apart relative to the spacing ofthe detection regions in the analysis unit. That is, for more accuratedetection, the photodiodes are placed apart at the same spacing as thespacing of the detection region.

The processor can be integrated into the same chip that contains theanalysis unit(s), or it can be separate, e.g., an independent microchipconnected to the analysis unit-containing chip via electronic leads thatconnect to the detection region(s) and/or to the discriminationregion(s), such as by a photodiode. The processor can be a computer ormicroprocessor, and is typically connected to a data storage unit, suchas computer memory, hard disk, or the like, and/or a data output unit,such as a display monitor, printer and/or plotter.

The types and numbers of molecules or cells, based on detection of areporter associated with or bound to the molecules or cells passingthrough the detection region, can be calculated or determined, and thedata obtained can be stored in the data storage unit. This informationcan then be further processed or routed to the data outlet unit forpresentation, e.g. histograms, of the types of molecules or cells (orlevels of a cell protein, saccharide), or some other characteristic. Thedata can also be presented in real time as the sample is flowing throughthe device.

With reference to FIG. 3B, the photodiode 78 can be larger in diameterthan the width of the main channel, forming a detection region 80 thatis longer (along the length of the main channel 82) than it is wide. Thevolume of such a detection region is approximately equal to thecross-sectional area of the channel above the diode multiplied by thediameter of the diode.

In a preferred sorting embodiment the detection region is connected bythe main channel to the discrimination region. The discrimination regionmay be located immediately downstream of the detection region, or may beseparated by a suitable length of channel. Constraints on the length ofchannel between the detection and discrimination regions are discussedbelow, with respect to the operation of the device. This length istypically between about 1 μm and about 2 cm. The discrimination regionis at the junction of the main channel and the branch channels. Itcomprises the physical location where molecules are directed into aselected branch channel. The means by which the molecules or cells aredirected into a selected branch channel may (i) be present in thediscrimination region, as in, e.g., electrophoretic or microvalve-baseddiscrimination, or (ii) be present at a distant location, as in, e.g.,electroosmotic or flow stoppage-based discrimination.

If desired, the device may contain a plurality of analysis units, i.e.,more than one detection and discrimination region, and a plurality ofbranch channels which are in fluid communication with and branch outfrom the discrimination regions. It will be appreciated that theposition and fate of molecules or cells in the discrimination region canbe monitored by additional detection regions installed, for example,immediately upstream of the discrimination region and/or within thebranch channels immediately downstream of the branch point. Theinformation obtained by the additional detection regions can be used bya processor to continuously revise estimates of the velocity of themolecules or cells in the channels and to confirm that molecules orcells having a selected characteristic enter the desired branch channel.

A group of manifolds (a region consisting of several channels which leadto or from a common channel) can be included to facilitate movement ofsample from the different analysis units, through the plurality ofbranch channels and to the appropriate solution outlet. Manifolds arepreferably microfabricated into the chip at different levels of depth.Thus, devices of the invention having a plurality of analysis units cancollect the solution from associated branch channels of each unit into amanifold, which routes the flow of solution to an outlet. The outlet canbe adapted for receiving, for example, a segment of tubing or a sampletube, such as a standard 1.5 ml centrifuge tube. Collection can also bedone using micropipettes.

6.3. Valve Structures

In an embodiment where pressurized flow is used, valves can be used toblock or unblock the pressurized flow of molecules or cells throughselected channels. A thin cantilever, for example, may be includedwithin a branch point, as shown in FIGS. 4A and 4B, such that it may bedisplaced towards one or the other wall of the main channel, typicallyby electrostatic attraction, thus closing off a selected branch channel.Electrodes are on the walls of the channel adjacent to the end of thecantilever. Suitable electrical contacts for applying a potential to thecantilever are also provided in a similar manner as the electrodes.Because the cantilever in FIG. 4B is parallel to the direction ofetching, it may be formed of a thin layer of silicon by incorporatingthe element into the original photoresist pattern. The cantilever ispreferably coated with a dielectric material such as silicon nitride, asdescribed in. Wise, et al., 1995 (46), for example, to prevent shortcircuiting between the conductive surfaces.

Alternatively, a valve may be situated within each branch channel,rather than at the branch point, to close off and terminate pressurizedflow through selected channels. Because the valves are locateddownstream of the discrimination region, the channels in this region maybe formed having a greater width than in the discrimination region,which simplifies the formation of valves.

A valve within a channel may be microfabricated, if desired, in the formof an electrostatically operated cantilever or diaphragm. Techniques forforming such elements are well known in the art (e.g., 24, 40, 46, 47,48). Typical processes include the use of selectively etched sacrificiallayers in a multilayer structure or, for example, the undercutting of alayer of silicon dioxide via anisotropic etching. For example, to form acantilever within a channel, as illustrated in FIGS. 4A and 4B, asacrificial layer 168 may be formed adjacent to a small section of anon-etchable material 170, using known photolithography methods, on thefloor of a channel, as shown in. FIG. 4A. Both layers can then be coatedwith, for example, silicon dioxide or another non-etchable layer, asshown at 172. Etching of the sacrificial layer deposits the cantilevermember 174 within the channel, as shown in FIG. 4B. Suitable materialsfor the sacrificial layer, non-etchable layers and etchant includeundoped silicon, p-doped silicon and silicon dioxide, and the etchantEDP (ethylene diamine/pyrocatechol), respectively. Because thecantilever in FIG. 4B is parallel to the direction of etching, it may beformed of a thin layer of silicon by incorporating the element into theoriginal photoresist pattern. The cantilever is preferably coated with adielectric material such as silicon nitride, as described in (46) forexample, to prevent short circuiting between the conductive surfaces.

The width of the cantilever or diaphragm should approximately equal thatof the channel, allowing for movement within the channel. If desired,the element may be coated with a more malleable material, such as ametal, to allow for a better seal. Such coating may also be employed torender a non-conductive material, such as silicon dioxide, conductive.

As above, suitable electrical contacts are provided for displacing thecantilever or diaphragm towards the opposing surface of the channel.When the upper surface is a glass cover plate, as described below,electrodes and contacts may be deposited onto the glass.

It will be apparent to one of skill in the field that other types ofvalves or switches can be designed and fabricated, using well knownphotolithographic or other microfabrication techniques, for controllingflow within the channels of the device. Multiple layers of channels canalso be prepared.

Operation of the valves or charging of the electrodes, in response tothe level of fluorescence measured from an analyte molecule, iscontrolled by the processor, which receives this information from thedetector. All of these components are operably connected in theapparatus, and electrical contacts are included as necessary, usingstandard microchip circuitry.

In preferred embodiments, an integrated semiconductor laser and/or anintegrated photodiode detector are included on the silicon wafer in thevicinity of the detection region. This design provides the advantages ofcompactness and a shorter optical path for exciting and/or emittedradiation, thus minimizing distortion.

The silicon substrate containing the microfabricated flow channels andother components is covered and sealed, preferably with a thin glass orquartz cover, although other clear or opaque cover materials may beused. when external radiation sources or detectors are employed, theinterrogation region is covered with a clear cover material to allowoptical access to the analyte molecules. Anodic bonding to a “PYREX”cover slip may be accomplished by washing both components in an aqueousH₂SO₄/H₂O₂ bath, rinsing in water, and then heating to about 350° C.while applying a voltage of, e.g., 45 OV.

6.4. Examples of Microchip Architecture for Sorting

As illustrated with respect to FIGS. 5A-5D, there are a number of waysin which cells can be routed or sorted into a selected branch channel.

FIG. 5A shows a discrimination region 102, which is suitable forelectrophoretic discrimination as the sorting technique. Thediscrimination region is preceded by a main channel 104. A junctiondivides the main channel into two branch channels 106 and 108. Thediscrimination region 102 includes electrodes 110 and 112, positioned onouter side walls of the branch channels 106 and 108, and which connectto leads 114 and 116. The leads are connected to a voltage source (notshown) incorporated into or controlled by a processor (not shown), asdescribed, infra. The distance (D) between the electrodes is preferablyless than the average distance separating the cells during flow throughthe main channel. The dimensions of the electrodes are typically thesame as the dimensions of the channels in which they are positioned, e.esuch that the electrodes are as high and wide as the channel.

The discrimination region shown in FIG. 5B is suitable for use in adevice that employs electro-osmotic flow, to move the molecules or cellsand bulk solution through the device. FIG. 4B shows a discriminationregion 122 which is preceded by a main channel 124. The main channelcontains a junction that divides the main channel into two branchchannels 126 and 128. An electrode 130 is placed downstream of thejunction of the main channel, for example near the sample inlet of mainchannel. Electrodes are also placed in each branch channel (electrodes132 and 134). The electrode 130 can be negative and electrodes 132 and134 can be positive (or vice versa) to establish bulk solution flowaccording to well-established principles of electro-osmotic flow (1E974: 25).

After a molecule or cell passes the detection region (not shown) andenters the discrimination region 122 (e.g. between the main channel andthe two branch channels) the voltage to one of the electrodes 132 or 134can be shut off, leaving a single attractive force that acts on thesolution and the molecule or cell to influence it into the selectedbranch channel. As above, the appropriate electrodes are activated afterthe molecule or cell has committed to the selected branch channel inorder to continue bulk flow through both channels. In one embodiment,the electrodes are charged to divert the flow into one branch channel,for example channel 126, which can be called a waste channel. Inresponse to a signal indicating that a molecule or cell has beenidentified or selected for collection, the charge on the electrodes canbe changed to divert the selected molecule or cell into the otherchannel (channel 128), which can be called a collection channel.

In another embodiment of the invention, shown in FIG. 5C, the moleculesor cells are directed into a predetermined branch channel via a valve140 in the discrimination region. The valve 140 comprises a thinextension of material to which a charge can be applied via an electrodelead 142. The valve 140 is shown with both channels open, and can bedeflected to close either branch channel by application of a voltageacross electrodes 144 and 146. A molecule or cell is detected and chosenfor sorting in the detection region. (not shown), and can be directed tothe appropriate channel by closing off the other channel, e.g. byapplying, removing or changing a voltage applied to the electrodes. Thevalve can also be configured to close one channel in the presence of avoltage, and to close the other channel in the absence of a voltage.

FIG. 5D shows another embodiment of a discrimination region of theinvention, which uses flow stoppage in one or more branch channels asthe discrimination means. The sample solution moves through the deviceby application of positive pressure at an end where the solution inletis located. Discrimination or routing of the molecules or cells isaffected by simply blocking a branch channel (145 or 148) or a branchchannel sample outlet using valves in a pressure-driven flow (147 or149). Due to the small size scale of the channels and theincompressibility of liquids, blocking the solution flow creates aneffective “plug” in the non-selected branch channel, thereby temporarilyrouting the molecule or cell together with the bulk solution flow intothe selected channel. Valve structures can be incorporated downstreamfrom the discrimination region, which are controlled by the detectionregion, as described herein.

Alternatively, the discrimination function represented in FIG. 5D may becontrolled by changing the hydrostatic pressure at the sample outlets ofone or both branch channels 145 or 148. If the branch channels in aparticular analysis unit have the same resistance to fluid flow, and thepressure at the sample inlet of the main channel of an analysis unit isP, then the fluid flow out of any selected branch channel can be stoppedby applying a pressure P/n at the sample outlet of the desired branchchannel, where n is the number of branch channels in the analysis unit.Accordingly, in an analysis unit having two branch channels, thepressure applied at the outlet of the branch to be blocked is P/2.

As shown in FIG. 5D, a valve is situated within each branch channel,rather than at the branch point, to close off and terminate pressurizedflow through selected channels. Because the valves are located at apoint downstream from the discrimination region, the channels in thisregion may be formed having a greater width than in the discriminationregion in order to simplify the formation of valves. The width of thecantilever or diaphragm should approximately equal the width of thechannel, allowing for movement within the channel. If desired, theelement may be coated with a more malleable material, such as a metal,to allow for a better seal. Such coating may also be employed to rendera non-conductive material, such as silicon dioxide, conductive. Asabove, suitable electrical contacts are provided for displacing thecantilever or diaphragm towards the opposing surface of the channel.When the upper surface is a glass cover plate, electrodes and contactsmay be deposited onto the glass.

6.5. Cascade Device

FIG. 6 shows a device with analysis units containing a cascade ofdetection and discrimination regions suitable for successive rounds ofpolynucleotide or cell sorting. Such a configuration may be used, forexample, with a polynucleotide or cells sorting device to generate aseries of samples containing “fractions” of polynucleotides, where eachfraction contains a specific size range of polynucleotide fragments(e.g., the first fraction contains 100-500 bp fragments, the next500-1000 bp fragments, and so on). In a cell sorting device, such acascade configuration may be used to sequentially assay the cell for,e.g., three different fluorescent dyes corresponding to expression ofthree different molecular markers. Samples collected at the outlets ofthe different branch channels contain pools of cells expressing definedlevels of each of the three markers. The number of reporters employed,and therefore the number of markers of interest, can be varied asdesired, e.g. to meet the needs of a particular experiment orapplication.

6.6. Microfabricated Polynucleotide Analysis Device

Also included in the present invention is a microfabricatedpolynucleotide analysis device suitable for quantitation and analysis ofthe size distribution of polynucleotide fragments in solution. Such adevice is a simplified version of the sorting device described above, inthat analysis units in the device need not contain a discriminationregion or branch channels, and the device need not contain a means fordirecting molecules to selected branch channels. Each analysis unitcomprises a single main channel containing a detection region asdescribed above. Since the optics which collect the optical signal(e.g., fluorescence) can be situated immediately adjacent the flowstream (e.g., diode embedded in the channel of a microscope objectiveadjacent a glass coverslip covering the channel), the signal-to-noiseratio of the signal collected using a microfabricated polynucleotideanalysis device of the invention is high relative to other types ofdevices. Specifically, the invention allows, e.g., the use ofoil-immersion high numerical aperture (N.A.) microscope objectives tocollect the light (e.g., 1.4 N.A.). Since the collection of light isproportional to the square of the N.A., a 1.4 N.A. objective providesabout a four-fold better signal than an 0.8 N.A. objective.

6.7. Microfabricated Cell Sorting Device

The invention also includes a microfabricated device for sortingreporter-labeled cells by the level of reporter they contain. The deviceis similar to polynucleotide-sorting devices described above, but isadapted for handling particles on the size scale of cells rather thanmolecules. This difference is manifested mainly in the dimensions of themicrofabricated channels, detection and discrimination regions.Specifically, the channels in the device are typically between about 20μm and about 500 μm in width and between about 20 μm and about 500 μm indepth, to allow for an orderly flow of cells in the channels. Similarly,the volume of the detection region in a cell sorting device is largerthan that of the polynucleotide sorting device, typically being in therange of between about 10 pl and 100 nl. To prevent the cells fromadhering to the sides of the channels, the channels (and coverslip)preferably contain a coating which minimizes cell adhesion. Such acoating may be intrinsic to the material from which the device ismanufactured, or it may be applied after the structural aspects of thechannels have been microfabricated. An exemplary coating has the surfaceproperties of a material such as “TEFLON”.

The device may be used to sort any procaryotic (e.g., bacterial) oreukaryotic (e.g., mammalian) cells which can be labeled (e.g., viaantibodies) with optically-detectable reporter molecules (e.g.,fluorescent dyes). Exemplary mammalian cells include human blood cells,such as human peripheral blood mononuclear cells (PBMCs). The cells canbe labeled with antibodies directed against any of a variety of cellmarker antigens (e.g., HLA DR, CD3, CD4, CD8, CD11a, CD11c, CD14, CD16,CD20, CD45, CD45RA, CD62L, etc.), and the antibodies can in turn bedetected using an optically-detectable reporter (either via directlyconjugated reporters or via labeled secondary antibodies) according tomethods known in the art.

It will be appreciated that the cell sorting device and method describedabove can be used simultaneously with multiple optically-detectablereporters having distinct optical properties. For example, thefluorescent dyes fluorescein (FITC), phycoerythrin (PE), and “CYCHROME”(Cy5-PE) can be used simultaneously due to their different excitationand emission spectra. The different dyes may be assayed, for example, atsuccessive detection and discrimination regions. Such regions may becascaded as shown in FIG. 6 to provide samples of cells having aselected amount of signal from each dye.

6.8. Microfabrication of a Silicon Device

Analytical devices having microscale flow channels, valves and otherelements can be designed and fabricated from a solid substrate material.Silicon is a preferred substrate material because of the well developedtechnology permitting its precise and efficient fabrication, but othermaterials may be used, including polymers such aspolytetrafluoroethylenes. Micromachining methods well known in the artinclude film deposition processes, such as spin coating and chemicalvapor deposition, laser fabrication or photolithographic techniques, oretching methods, which may be performed by either wet chemical or plasmaprocesses. (See, for example, Angell et al. (48) and Manz et al. (49).

FIGS. 7A-7D illustrate the initial steps in microfabricating thediscrimination region portion of a nucleic acid sorting device (e.g.Device 20 in FIG. 1) by photolithographic techniques. As shown, thestructure includes a silicon substrate 160. The silicon wafer whichforms the substrate is typically washed in a 4:1H₂SO₄/H₂O bath, rinsedin water and spun dry. A layer 162 of silicon dioxide, preferably about0.5 μm in thickness, is formed on the silicon, typically by heating thesilicon wafer to 800-1200° C. in an atmosphere of steam. The oxide layeris then coated with a photoresist layer 164, preferably about 1 μminch-thickness. Suitable negative or positive resist materials are wellknown. Common negative resist materials include two-componentbisarylazide/rubber resists. Positive resist materials includepolymethyl-methacrylate (PMMA) and two component diazoquinone/phenolicresin materials. See, e.g., “Introduction to microlithography”, Thompson(47).

The coated laminate is irradiated through a photomask 166 imprinted witha pattern corresponding in size and layout to the desired pattern of themicrochannel. Methods for forming photomask having desired photomaskpatterns are well known. For example, the mask can be prepared byprinting the desired layout on an overhead transparency using a highresolution (3000 dpi) printer. Exposure is carried out on standardequipment such as a Karl Sass contact lithography machine.

In the method illustrated in FIGS. 7A-5D, the photoresist is a negativeresist, meaning that exposure of the resist to a selected wavelength,e.g., UV, light produces a chemical change that renders the exposedresist material resistant to the subsequent etching step. Treatment witha suitable etchant removes the unexposed areas of the resist, leaving apattern of bare and resist-coated silicon oxide on the wafer surface,corresponding to the layout and dimensions of the desired microstructures. In this example, because a negative resist was used, thebare areas correspond to the printed layout on the photomask. The waferis now treated with a second etchant material, such as a reactive ionetch. (R.E.), effective to dissolve the exposed areas of silicondioxide. The remaining resist is removed, typically with hot aqueousH₂SO₄. The remaining pattern of silicon dioxide (162) now serves as amask for the silicon (160). The channels are etched in the unmaskedareas of the silicon substrate by treating with a KO etching solution.Depth of etching is controlled by time of treatment. Additionalmicrocomponents may also be formed within the channels by furtherphotolithography and etching steps, as discussed below.

Depending on the method to be used for directing the flow of moleculesthrough the device, electrodes and/or valves are fabricated into theflow channels. A number of different techniques are available forapplying thin metal coatings to a substrate in a desired pattern. Theseare reviewed in, for example, Krutenat, Kirk-Othmer 3rd ed., Vol. 15,pp. 241-274 (43), incorporated herein by reference. A convenient andcommon technique used in fabrication of microelectronic circuitry isvacuum deposition. For example, metal electrodes or contacts may beevaporated onto a substrate using vacuum deposition and a contact maskmade from, e.g., a “MYLAR” sheet. Various metals such as platinum, gold,silver or indium/tin oxide (ITO) may be used for the electrodes.

Deposition techniques allowing precise control of the area of depositionare preferred for application of electrodes to the side walls of thechannels in the device. Such techniques are described, for example, inKrutenat (43), above, and references cited therein. They include plasmaspraying, where a plasma gun accelerates molten metal particles in acarrier gas towards the substrate, and physical vapor deposition usingan electron beam, where atoms are delivered on line-of-sight to thesubstrate from a virtual point source. In laser coating, a laser isfocused onto the target point on the substrate, and a carrier gasprojects powdered coating material into the beam, so that the moltenparticles are accelerated toward the substrate.

Another technique allowing precise targeting uses an electron beam toinduce selective decomposition of a previously deposited substance, suchas a metal salt, to a metal. This technique has been used to producesub-micron circuit paths (e.g., 37).

6.9. Elastomeric Microfabricated Device

This Example demonstrates the manufacture of a disposablemicrofabricated device, which can function as a stand-alone device or asa component of an integrated microanalytical chip, in sorting moleculesor cells. In particular, this example describes exemplary microfluidicdevices that are manufactured from an elastomer material (e.g., asilicone elastomer). Other elastomer materials that may be used includesilicone elastomers such as polydimethylsiloxane (PDMS) (see, e.g.,Subsection 6.12.1, infra). Such materials are particularly preferred inembodiments, e.g., wherein the features of a microfluidic device (e.g.,channel widths and depths, valves and pumps, etc.) have sizes thatapproach or are below the limits of optical diffraction and aretherefore smaller than can be obtained through traditional opticallithography techniques.

Micrometer or nanometer scale microfluidic devices may be readilymicrofabricated with such materials, e.g., using the replica molding or“soft lithography” techniques described herein and by Xia and Whitesides(24). However, other microfabrication techniques are also known in theart and may be used to fabricate microfluidic devices of this invention;e.g., “nano-imprint lithography” techniques (86, 87). Additionalmaterials and methods that can be used to manufacture microfluidicdevices of this invention are disclosed in U.S. Provisional PatentApplication Ser. No. 60/249,362, filed Nov. 16, 2000 and incorporatedherein, by reference, in its entirety.

6.9.1. Preparation of the Microfabricated Device

A silicon wafer was etched and fabricated as described above and in(27). After standard contact photolithography techniques to pattern theoxide surface of the silicon wafer, a C₂F₂/CHF₃ gas mixture was used toetch the wafer by R.E. The silicon wafer was then subjected to furtheretch with KO to expose the silicon underneath the oxide layer, therebyforming a mold for the silicone elastomer. The silicon mold forms a “T”arrangement of channels. The dimensions of the channels may rangebroadly, having approximately 5×4 μm dimension.

The etching process is shown schematically in FIG. 8. Standardmicromachining techniques were used to create a negative master mold outof a silicon wafer. The disposable silicone elastomer chip was made bymixing General Electric RTV 615 components (32) together and pouringonto the etched silicon wafer. After curing in an oven for 2 hours at80° C., the elastomer was peeled from the wafer and bonded hermeticallyto a glass cover slip for sorting. To make the elastomer hydrophilic theelastomer chip was immersed in HCl (pH=2.7) at 60 degrees C. for 40 to60 min. Alternatively, the surface could have been coated withpolyurethane (3% w/v in 95% ethanol and diluted 10× in ethanol). It isnoted that the master wafer can be reused indefinitely. The device shownhas channels that are 100 μm wide at the wells, narrowing to 3 μm at thesorting junction (discrimination region). The channel depth is 4 μm, andthe wells are 2 mm in diameter. These dimensions can be modifiedaccording to the size range of the molecules or cells to be analyzed orsorted.

6.9.2. Detection Apparatus

In this embodiment the device was mounted on an inverted opticalmicroscope (Zeiss Axiovert 35) as shown in FIG. 9. In this system, theflow control can be provided by voltage electrodes for electro-osmoticcontrol or by capillaries for pressure-driven control. The detectionsystem can be photomultiplier tubes or photodiodes, depending upon theapplication. The inlet well and two collection wells were incorporatedinto the elastomer chip on three sides of the “T” forming three channels(FIG. 7). The chip was adhered to a glass coverslip and mounted onto themicroscope.

6.10. Operation of a Microfabricated Polynucleotide Analysis Device

The operation of a polynucleotide analysis chip is described. Thisexample refers to polynucleotides, but it will be appreciated that othermolecules may be analyzed or sorted using similar methods and devices.Likewise, cells can be processed using similar methods and devices,adapted to the appropriate size.

A solution of reporter-labeled polynucleotides is prepared as describedbelow and introduced into the sample inlet end(s) of the analysisunit(s). The solution may be conveniently introduced into a reservoir,such as reservoir 48 of FIG. 1, via a port or connector, such asconnector 70 in FIG. 2, adapted for attachment to a segment of tubing,such as liquid chromatography or HPLC tubing.

It is typically advantageous to “hydrate” the device (i.e., fill thechannels of the device with the solvent, e.g., water or a buffersolution, in which the polynucleotides will be suspended) prior tointroducing the polynucleotide-containing solution. Such hydrating canbe achieved by supplying water or the buffer solution to the devicereservoir and applying hydrostatic pressure to force the fluid throughthe analysis unit(s).

Following such hydration, the polynucleotide-containing solution isintroduced into the sample inlets of the analysis unit(s) of the device.As the stream of labeled polynucleotides (e.g., tagged with a reportersuch as a fluorescent dye) is passed in a single file manner through thedetection region, the optical signal (e.g., fluorescence) from theoptically-detectable reporter moieties on each molecule are quantitatedby an optical detector and converted into a number used in calculatingthe approximate length of polynucleotide in the detection region.

Exemplary reporter moieties, described below in reference to samplepreparation, include fluorescent moieties which can be excited to emitlight of characteristic wavelengths by an excitation light source.Fluorescent moieties have an advantage in that each molecule can emit alarge number of photons (e.g., upward of 106) in response to excitingradiation. Suitable light sources include lasers, laser diodes,high-intensity lamps, e.g., mercury lamps, and the like. In embodimentswhere a lamp is used, the channels are preferably shielded from thelight in all regions except the detection region, to avoid bleaching ofthe label. In embodiments where a laser is used, the laser can be set toscan across a set of detection regions from different analysis units.Other optically-detectable reporter moieties include chemiluminescentmoieties, which can be used without an excitation light source.

Where laser diodes are used as a light source, the diodes may bemicrofabricated into the same chip that contains the analysis units.Alternatively, the laser diodes may be incorporated into a second chip(laser diode chip; LDC) that is placed adjacent to the chip such thatthe laser light from the diodes shines on the detection regions. Thephotodiodes in the LDC are preferably placed at a spacing thatcorresponds to the spacing of the detection regions in the chip.

The level of reporter signal is measured using an optical detector, suchas a photodiode (e.g., an avalanche photodiode), a fiber-optic lightguide leading, e.g., to a photomultiplier tube, a microscope with a highnumerical aperture (N.A.) objective and an intensified video camera,such as a SIT camera, or the like. The detector may be microfabricatedor placed into the PAC itself (e.g., a photodiode as illustrated inFIGS. 3A and 3B), or it may be a separate element, such as a microscopeobjective.

In cases where the optical detector is a separate element, it isgenerally necessary to restrict the collection of signal from thedetection region of a single analysis unit. It may also be advantageousto scan or move the detector relative to the polynucleotide analysisunit (“PAC”), preferably by automation. For example, the PAC can besecured in a movable mount (e.g., a motorized/computer-controlledmicromanipulator) and scanned under the objective. A fluorescencemicroscope, which has the advantage of a built-in excitation lightsource (epifluorescence), is preferably employed for detection of afluorescent reporter.

Since current microfabrication technology enables the creation ofsub-micron structures employing the elements described herein, thedimensions of the detection region are influenced primarily by the sizeof the molecules under study. These molecules can be rather large bymolecular standards. For example, lambda DNA (˜50 kb) in solution has adiameter of approximately 0.5 μm. Accordingly, detection regions usedfor detecting polynucleotides in this size range have a cross-sectionalarea large enough to allow such a molecule to pass through without beingsubstantially slowed down relative to the flow of the solution carryingit and causing a “bottle neck”. The dimensions of a channel should thusbe at least about twice, preferably at least about five times as largeper side or in diameter as the diameter of the largest molecule thatwill be passing through it.

Another factor important to consider in the practice of the invention isthe optimal concentration of polynucleotides in the sample solution,particularly in embodiments which process or analyze one molecule orcell at a time. The concentration, should be dilute enough so that alarge majority of the polynucleotide molecules pass through thedetection region one by one, with only a small statistical chance thattwo or more molecules pass through the region simultaneously. This is toinsure that for the large majority of measurements, the level ofreporter measured in the detection region corresponds to a singlemolecule, rather than two individual molecules.

The parameters which govern this relationship are the volume of thedetection region and the concentration of molecules in the samplesolution. The probability that the detection region will contain two ormore molecules (P_(≧2)) can be expressed as

P _(≧2)=1−{1+[DNA]*V}*e ^(−[DNA]*V)

where [DNA] is the concentration of polynucleotides in units ofmolecules per μm³ and V is the volume of the detection region in unitsof μm³.

It will be appreciated that P_(>2) can be minimized by decreasing theconcentration of polynucleotides in the sample solution. However,decreasing the concentration of polynucleotides in the sample solutionalso results in increased volume of solution processed through thedevice and can result in longer run times. Accordingly, the objectivesof minimizing the simultaneous presence of multiple molecules in thedetection chamber (to increase the accuracy of the sorting) needs to bebalanced with the objective of generating a sorted sample in areasonable time in a reasonable volume containing an acceptableconcentration of polynucleotide molecules.

The maximum tolerable P_(≧2) depends on the desired “purity” of thesorted sample. The “purity” in this case refers to the fraction ofsorted polynucleotides that are in the specified size range, and isinversely proportional to P_(≧2). For example, in applications wherehigh purity is not required, such as the purification of a particularrestriction fragment from an enzymatic digest of a portion of vectorDNA, a relatively high P_(≧2) (e.g., P_(≧2)=0.2) may be acceptable. Formost applications, maintaining P_(≧2) at or below about 0.1 providessatisfactory results.

In an example where P_(≧2) is equal 0.1, it is expected that in about10% of measurements, the signal from the detection region will be due tothe presence of two or more polynucleotide molecules. If the totalsignal from these molecules is in the range corresponding to the desiredsize fragment, these (smaller) molecules will be sorted into the channelor tube containing the desired size fragments.

The DNA concentration needed to achieve a particular value P_(≧2) in aparticular detection volume can be calculated from the above equation.For example, a detection region in the shape of a cube 1 μm³ per sidehas a volume of 1 femtoliter (fl). A concentration of moleculesresulting, on average, in one molecule per fl, is about 1.7 nM. Using aP_(≧2) of about 0.1, the polynucleotide concentration in a sampleanalyzed or processed using such a 1 fl detection region volume isapproximately 0.85 nM, or roughly one DNA molecule per 2 detectionvolumes ([DNA]*V=˜0.5). If the concentration of DNA is such that [DNA]*Vis 0.1, P_(≧2) is less than 0.005; i.e., there is less than a one halfof one percent chance that the detection region will at any given timecontain two of more fragments.

The signal from the optical detector is routed, e.g., via electricaltraces and pins on the chip, to a processor, which processes the signalsinto values corresponding to the length of the polynucleotide givingrise to the signal. These values are then compared, by the processor, topre-loaded instructions containing information on which branch channelmolecules of a particular size range will be routed into. Following adelay period that allows the molecule from which the reporter signaloriginated to arrive at the discrimination region, the processor sends asignal to actuate the active elements in the discrimination region suchthat the molecule is routed into the appropriate branch channel.

The delay period is determined by the rate at which the molecules movethrough the channel (their velocity relative to the walls of thechannel) and the length of the channel between the detection region andthe discrimination region. In cases where the sample solution is movedthrough the device using hydrostatic pressure (applied, e.g., aspressure at the inlet end and/or suction at the outlet end), thevelocity is typically the flow rate of the solution. In cases where themolecules are pulled through the device using some other means, such asvia electro-osmotic flow with an electric field set up between the inletend and the outlet end, the velocity as a function of molecule size canbe determined empirically by running standards, and the velocity for aspecific molecule calculated based on the size calculated for it fromthe reporter signal measurement.

A relevant consideration with respect to the velocity at which thepolynucleotide molecules move through the device is the shear force thatthey may be subject to. At the channel dimensions contemplated herein,the flow through the channels of the device is primarily laminar flowwith an approximately parabolic velocity profile. Since thecross-sectional area of the channels in the device can be on the sameorder of magnitude as the diameter of the molecules being analyzed,situations may arise where a portion of a particular molecule is verynear the wall of the channel, and is therefore in a low-velocity region,while another portion of the molecule is near the center of the channel,i.e., in a high-velocity region. This situation creates a shear force(F) on the molecule, which can be estimated using the followingexpression:

F=6πηR _(λ) V

where R_(λ) is the radius of the molecule and η is the viscosity of thesolution. This expression assumes that the molecule is immobilized on astationary surface and subject to uniform flow of velocity V.

The amount of force necessary to break a double stranded fragment of DNAis approximately 100 pN. Accordingly, the maximal shear force that themolecules are subjected to should preferably be kept below this value.Substituting appropriate values for the variables in the aboveexpression for lambda DNA yields a maximum velocity of about 1 cm/secfor a channel 1 μm in radius (i.e., a channel of a dimension where oneportion of the lambda molecule can be at or near the wall of the channelwith the opposite side in the center of the channel). Since devicesdesigned for use with such large molecules will typically have channelsthat are considerably larger in diameter, the maximum “safe” velocitywill typically be greater than 1 cm/sec.

As discussed above, the sample solution introduced into a device of theinvention should be dilute enough such that there is a high likelihoodthat only a single molecule occupies the detection region at any giventime. It follows then that as the solution flows through the devicebetween the detection and discrimination regions, the molecules will bein “single file” separated by stretches of polynucleotide-free solution.The length of the channel between the detection and discriminationregion should therefore not be so long as to allow random thermaldiffusion to substantially alter the spacing between the molecules. Inparticular, the length should be short enough that it can be traversedin a time short enough such that even the smallest molecules beinganalyzed will typically not be able to diffuse and “switch places” inthe line of molecules.

The diffusion constant of a 1 kb molecule is approximately 5 μm²/sec;the diffusion equation gives the distance that the molecule diffuses intime t as

<X ² >˜Dt

Using this relationship, it can be appreciated that a 1 kbp fragmenttakes about 0.2 seconds to diffuse 1 μm. The average spacing ofmolecules in the channel is a function of the cross-sectional area ofthe channel and the molecule concentration, the latter being typicallydetermined in view of acceptable values of P_(≧2) (see above). From theabove relationships, it is then straightforward to calculate the maximumchannel length between the detection and discrimination region whichwould ensure that molecules don't “switch places”. In practice, thechannel length between the detection and discrimination regions isbetween about 1 μm and about 2 cm.

As illustrated above with respect to FIGS. 5A-D, there are a number ofways in which molecules can be routed or sorted into a selected branchchannel. For example, in a device employing the discrimination regionshown in FIG. 4A, the solution is preferably moved through the device byhydrostatic pressure. Absent any field applied across electrodes 110 and112, a molecule would have an equal probability of entering one or theother of the two branch channels 106 and 108. The sorting isaccomplished by the processor temporarily activating a voltage sourceconnected to the electrode leads 114 and 116 just before or at the timethe molecule to be routed enters the junction of the main channel andthe two branch channels. The resulting electric field exerts a force onthe negatively-charged DNA molecule biasing it toward thepositively-charged electrode. The molecule will then be carried down thebranch channel containing the positively-charged electrode by the bulksolution flow. The electric field is turned off when the molecule hascommitted itself to the selected channel. As soon as the molecule clearsthe corner from the discrimination region and into the branch channel,it escapes effects of the electric field that will be applied to thenext molecule in the solution stream.

The discrimination region shown in FIG. 5B is designed for use in adevice that employs electroosmotic flow, rather than flow induced byhydrostatic pressure, to move both the polynucleotides and bulk solutionthrough the device. Electrodes are set up in the channels at the inletand outlet ends of the device. Application of an electric field at theends of the channels (with electrode 130 being negative, and electrodes132 and 134 being positive) sets up bulk solution flow according towell-established principles of electroosmotic flow (see, e.g., 36). Whena specific polynucleotide molecule enters the junction region betweenthe main channel and the two branch channels, the voltage to one ofeither electrodes 132 or 134 is shut off, leaving a single attractiveforce, acting on the solution and the DNA molecule, into the selectedbranch channel. As above, both branch channel electrodes are activatedafter the molecule has committed to the selected branch channel in orderto continue bulk flow through both channels.

In another embodiment of the invention the polynucleotides are directedinto a selected branch channel via a valve in the discrimination region.An exemplary valve is shown in FIG. 5C. The valve consists of a thinextension of material 140 which can be charged via an electrode 142. Theextension can then be deflected to close one or the other of the branchchannels by application of an appropriate voltage across electrodes 144and 146. As above, once the molecule has committed, the voltage can beturned off.

In a device in which the sample solution is moved through the device byapplication of positive pressure at the sample inlet end(s) of theanalysis unit(s), the discrimination function may be affected by simplyblocking branch channel sample outlets into which the sample is notsupposed to go, and leaving the selected outlet open. Due to the smallsize scale of the channels and the incompressibility of liquids,blocking the solution flow creates an effective “plug” in the unselectedbranch channels, routing the molecule along with the bulk solution flowinto the selected channel. This embodiment is illustrated in FIG. 4D. Itcan be achieved by, for example, incorporating valve structuresdownstream of the discrimination region.

Alternatively, the discrimination function may be affected by changingthe hydrostatic pressure at the sample outlets of the branch channelsinto which the sample is not supposed to go. Specifically, if the branchchannels in a particular analysis unit all offer the same resistance tofluid flow, and the pressure at the sample inlet of the main channel ofan analysis unit is P, then the fluid flow out of any selected branchchannel can be stopped by applying a pressure P/n at the sample outletof that branch channel, where n is the number of branch channels in thatanalysis unit. Accordingly, in an analysis unit having two branchchannels, the pressure applied at the outlet of the branch to be blockedis P/2.

It will be appreciated that the position and fate of the molecules inthe discrimination region can be monitored by additional detectionregions installed, e.g., immediately upstream of the discriminationregion and/or in the branch channels immediately downstream of thebranch point. This information be used by the processor to continuouslyrevise estimates of the velocity of the molecules in the channels and toconfirm that molecules having selected size characteristics end up inthe selected branch channel.

Solution from the branch channels is collected at the outlet ends of theanalysis units. As described above, devices with a plurality of analysisunits typically collect the solution from corresponding branch channelsof each unit into a manifold, which routes the solution flow to anoutlet port, which can be adapted for receiving, e.g., a segment oftubing or a sample tube, such as a standard 1.5 ml centrifuge tube.

The time required to isolate a desired quantity of polynucleotidedepends on a number of factors, including the size of thepolynucleotide, the rate at which each analysis unit can process theindividual fragments, and the number of analysis units per chip-, andcan be easily calculated using basic formulas. For example, a chipcontaining 1000 analysis units, each of which can sort 1000 fragmentsper second, could isolate 0.1 μg of 10 kb DNA in about 2.5 hours.

6.11. Other Microfabricated Devices of the Invention

Operation of a microfabricated cell sorting device is essentially asdescribed above with respect to the polynucleotide sorting device. Sincecells typically do not have predictable a net charge, the directingmeans are preferably ones employing a valve in the discrimination regionas described above, or flow stoppage, either by valve or hydrostaticpressure.

Operation of a microfabricated analysis device is accomplishedessentially as is described above, except that functions relating tosorting polynucleotide molecules into branch channels don't need to beperformed. The processor of such analysis devices is typically connectedto a data storage unit, such as computer memory, hard disk or the like,as well as to a data output unit, such as a display monitor, printerand/or plotter. The sizes of the polynucleotide molecules passingthrough the detection region are calculated and stored in the datastorage unit. This information can then be further processed and/orrouted to the data output unit for presentation as, e.g., histograms ofthe size distribution of DNA molecules in the sample. The data can, ofcourse, be presented in real time as the sample is flowing through thedevice, allowing the practitioner of the invention to continue the runonly as long as is necessary to obtain the desired information.

In preferred molecular (e.g. DNA, polynucleotide or polypeptide)analysis and sorting embodiments, a microfabricated chip of theinvention has a detection volume of about 10 to about 5000 femtoliters(fl), preferably about 50 to about 1000 fl, and most preferably on theorder of about 200 fl. In preferred cell analysis and sortingembodiments, a microfabricated chip of the invention has a detectionvolume of approximately 1 to 1,000,000 femtoliters (fl), preferablyabout 200 to 500 fl, and most preferably about 375 fl.

6.12. Exemplary Embodiment and Additional Parameters 6.12.1.Microfluidic Chip Fabrication

In a preferred embodiment, the invention provides a “T” on “Y” shapedseries of channels molded into optically transparent silicone rubber orPolyDiMethylSiloxane (PDMS), preferably PDMS. This is cast from a moldmade by etching the negative image of these channels into the same typeof crystalline silicon wafer used in semiconductor fabrication. Asdescribed above, the same techniques for patterning semiconductorfeatures are used to form the pattern of the channels. The uncuredliquid silicone rubber is poured onto these molds placed in the bottomof a Petri dish. To speed the curing, these poured molds are baked.After the PDMS has cured, it is removed from on top of the mold andtrimmed. In a chip with one set of channels forming a “T”, three holesare cut into the silicone rubber at the ends of the “T”, for exampleusing a hole cutter similar to that used for cutting holes in cork, andsometimes called cork borers. These holes form the sample, waste andcollection wells in the completed device. In this example, the hole atthe bottom end of the T is used to load the sample. The hole at one armof the T is used for collecting the sorted sample while the opposite armis treated as waste. Before use, the PDMS device is placed in a hot bathof HCl to make the surface hydrophilic. The device is then placed onto aNo. 1 (150 μm thick) (25×25 mm) square microscope cover slip. The coverslip forms the floor (or the roof) for all three channels and wells. Thechip has a detection region as described above.

Note that any of or all of these manufacturing and preparation steps canbe done by hand, or they can be automated, as can the operation and useof the device.

The above assembly is placed on an inverted Zeiss microscope. A carrierholds the cover slip so that it can be manipulated by the microscope'sx-y positioning mechanism. This carrier also has mounting surfaces whichsupport three electrodes, which implement the electro-osmotic and/orelectrophoretic manipulation of the cells or particles to be analyzedand sorted. The electrodes are lengths of platinum wire taped onto asmall piece of glass cut from a microscope slide. The wire is bent intoa hook shape, which allows it to reach into one of the wells from above.The cut glass acts as a support platform for each of the electrodes.They are attached to the custom carrier with double-sided tape. Thisallows flexible positioning of the electrodes. Platinum wire ispreferred for its low rate of consumption (long life) in electrophoreticand electro-osmotic applications, although other metals such as goldwire may also be used.

6.12.2. Device Loading

To operate the device for sorting, one of the wells, e.g. the collectionor waste well, is first filled with buffer. All three channels, startingwith the channel connected to the filled well, wick in buffer viacapillary action and gravity. Preferably, no other well is loaded untilall the channels fill with buffer, to avoid the formation of airpockets. After the channels fill the remaining wells can be loaded withbuffer, as needed, to fill or equilibrate the device. The input orsample well is typically loaded last so that the flow of liquid in thechannels is initially directed towards it. Generally, equal volumes ofbuffer or sample are loaded into each well. This is done in order toprevent a net flow of liquid in any direction once all of the wells areloaded, including loading the sample well with sample. In thisembodiment, it is preferred that the flow of material through the device(i.e. the flow of sample) be driven only by the electrodes, e.g. usingelectro-osmotic and/or electrophoretic forces. The electrodes may be inplace during loading, or they can be placed into the wells afterloading, to contact the buffer.

6.12.3. Electrodes

Two of the above electrodes are driven by high voltage operationalamplifiers (op-amps) capable of supplying voltages of +−150 V. The thirdelectrode is connected to the electrical ground (or zero volts) of thehigh voltage op-amp electronics. For sorting operation the drivenelectrodes are placed in the collection and waste wells. The groundelectrode is placed in the sample well. The op-amps amplify, by a factorof 30, a control voltage generated by two digital to analog converters(DACs). The maximum voltage these DACs generate is +−5 V, whichdetermines the amplification factor of 30. The 150 V limit is determinedby the power supply to the amplifiers, which are rated for +−175 V.These DACs reside on a card (a Lab PC 1200 Card, available from NationalInstruments, Austin, Tex.) mounted in a personal computer. The card alsocontains multiple channels of analog to digital converters (ADC's) oneof which is used for measuring the signal generated by photomultipliertubes (PMTs). This card contains two DACs. A third DAC can be used todrive the third electrode with an additional high voltage op amp. Thiswould provide a larger voltage gradient, if desired, and some additionaloperational flexibility.

Without being bound by any theory, it is believed that the electrodesdrive the flow of sample through the device using electro-osmotic orelectrophoretic forces, or both. To start the movement of molecules,cells or particles to be sorted, a voltage gradient is established inthe channels. This is done by generating a voltage difference betweenelectrodes.

In this example, the voltage of the two driven electrodes is raised orlowered with respect to the grounded electrode. The voltage polaritydepends on the charge of the molecules, cells or particles to be sorted(if they are charged), on the ions in the buffer, and on the desireddirection of flow. Because the electrode at the sample well in theexamples is always at zero volts with respect to the other twoelectrodes, the voltage at the “T” intersection or branch point will beat a voltage above or below zero volts, whenever the voltage of theother two electrodes is raised or lowered. Typically, the voltage is setor optimized, usually empirically, to produce a flow from the samplewell, toward a downstream junction or branch point where two or morechannels meet In this example, where two channels are used, one channelis typically a waste channel and terminates in a waste well, the otherchannel is a collection channel and terminates in a collection well.

To direct the molecules, particles or cells to a particular channel orarm of the “T” (e.g. collection or waste), the voltage at the electrodein one well (or multiple wells, in multi-channel embodiments) is madethe same as the voltage at the junction or branch point, where thechannels meet. The voltage of the electrode at one well of the two ormore wells is raised or lowered, to produce a gradient between that welland the branch point. This causes the flow to move down the channeltowards and into the well, in the direction produced by the gradient.Typically, the voltage of the electrode at the waste well is raised orlowered with respect to the voltage at the collecting well, to directthe flow into the waste channel and the waste well, until a molecule,particle or cell is identified for collection. The flow is diverted intothe collection channel and collection well by adjusting the voltages atthe electrodes to eliminate or reduce the gradient toward the wastewell, and provide or increase the gradient toward the collection well.For example, in response to a signal indicating that a molecule or cellhas been detected for sorting, by examination in a detection regionupstream of the branch point, the voltage at the waste and collectionpoints can be switched, to divert the flow from one channel and well tothe other.

The voltage at the branch point (the intersection voltage) is determinedby the voltage gradient desired (e.g. Volts/mm) times the distance fromthe sample well electrode to the branch point (gradient×distance), whichin this example is placed where all of the channels of the “T”intersect. The gradient also determines the voltage at the waste orcollection electrode (gradient×distance from sample well to collectionwell).

Conceptually, the channels and wells of the “T” can be treated as anetwork of three resistors. Each segment of the “T” behaves as aresistor whose resistance is determined by the conductivity of thebuffer and the dimensions of the channel. A voltage difference isprovided across two of the resistors, but not the third. If theelectrodes in each of the three wells is equidistant from the branchpoint, then each channel will have the same resistance.

For example, assume that each section of the “T” has 100 K ohms ofresistance. If 100 volts is applied across two of the resistors and thethird resistor is left unconnected, the current at the junction of thetwo resistors would be 50 volts. If a voltage source of 50 volts isconnected to the end of the third resistor, the voltage at the junctiondoes not change. That is, a net of zero volts is established across thethird resistor; there is no voltage gradient and a flow is not initiatedor changed. If a different voltage is applied, a gradient can beestablished to initiate or direct the flow into one channel or another.For example, to change the direction of flow from one arm of the “T” tothe other, the voltage values of the two driven electrodes are swapped.The junction voltage remains the same. If the electrode distances fromthe “T” intersection are not equal, then the voltages can be adjusted toaccommodate the resulting differences in the effective channelresistance. The end result is still the same. The electrode in the wellof the channel which is temporarily designated not to receive particlesor cells is set at the voltage of the “T” intersection. The voltage atthe other driven electrode is set to provide a gradient that directsmolecule, cell or particle flow into that well. Thus, cells or particlescan be sent down one channel or another, and ultimately into one well oranother, by effectively opening one channel with a net or relativevoltage gradient while keeping the other channel or channels closed by anet or relative voltage gradient of zero.

In a preferred embodiment for sorting according to the invention, aslight flow down the channel that is turned “off” is desired. This keepsthe molecules or cells moving away from the branch point (the “T”junction), particularly those which have already been directed to thatchannel. Thus, a small non-zero gradient is preferably established inthe “off” or unselected channel. The selected channel is provided with asignificantly higher gradient, to quickly and effectively divert thedesired molecules or cells into that channel.

The placement of the wells and their electrodes with respect to thebranch point, and in particular their distance from the branch point, isan important factor in driving the flow of molecules or cells asdesired. As the wells and electrodes are brought closer to the branchpoint, it becomes more important to precisely place the electrodes, orprecisely adjust the voltages.

6.12.4. Detection Optics

In this example, a Ziess Axiovert 35 inverted microscope is used fordetection of molecules or cells for sorting. The objective lens of thismicroscope faces up, and is directed at the detection region of thedescribed microfluidic chip, through the coverslip which in this exampleis the floor of the device. This microscope contains all the componentsfor epifluorescence microscopy. See, Inoue pp 67-70, 91-97 (63). In thisembodiment a mercury arc lamp or argon ion laser is used as the lightsource. The mercury lamp provides a broad spectrum of light that canexcite many different fluorophores. The argon ion laser has greaterintensity, which improves the detection sensitivity but is generallyrestricted to fluorophores that excite at 488 or 514 nm. The mercurylamp is used, for example, to sort beads as described elsewhere herein.The laser is used for sorting GFP E. coli bacterial cells as describedelsewhere herein. The high power argon ion beam is expanded to fill theillumination port of the microscope, which matches the opticalcharacteristics of the mercury arc lamp and provides a fairly uniformillumination of the entire image area in a manner similar to the mercurylamp. However, it is somewhat wasteful of the laser light. If a lowerpowered laser is used, the laser light is focused down to coincide withthe detection region of the chip, to achieve the same or similarillumination intensity and uniformity with less power consumption.

The objective used in the example is an Olympus PlanApo 60×1.4 N.A. oilimmersion lens. The optics are of the infinity corrected type. An oilimmersion lens enables collecting a substantial percentage of the 180degree hemisphere of emitted light from the sample. This enables some ofthe highest sensitivity possible in fluorescence detection. Thismicroscope has 4 optical ports including the ocular view port. Eachport, except the ocular, taps −20% of the available light collected fromthe sample when switched into the optical path. Only the ocular port canview 100% of the light collected by the objective. In this embodiment, acolor video camera is mounted on one port, another has a Zeissadjustable slit whose total light output is measured with aphotomultiplier tube (PMT). The fourth port is not used.

The microscope focuses the image of the sample onto the plane of theadjustable slit. An achromatic lens collimates the light from the slitimage onto the active area of the PMT. Immediately in front of the PMTwindow an optical band pass filter is placed specific to thefluorescence to be detected. The PMT is a side on-type and does not havea highly uniform sensitivity across its active area. By relaying theimage to the PMT with the achromat, this non-uniformity is averaged andits effect on the measured signal is greatly minimized. This alsoenables near ideal performance of the bandpass filter. A 20% beamsplitter has been placed in the light path between the achromat andfilter. An ocular with a reticle re-images this portion of thecollimated light. This enables viewing the adjustable slit directly, toinsure that the detection area that the PMT measures is in focus andaligned. The adjustable slit allows windowing a specific area of thechannel for detection. Its width, height, and x, y position areadjustable, and conceptually define a detection region on the chip. Inthis embodiment, the microscope is typically set to view a 5 μm (micron)length of the channel directly below the “T” intersection.

The PMT is a current output device. The current is proportional to theamount of light incident on the photocathode. A transimpedance amplifierconverts this photo-current to a voltage that is digitized by the Lab PC1200 card. This allows for interpreting the image to select cells orparticles having an optical reporter for sorting, as they pass throughthe detection region, based for example on the amount of light orfluorescence measured as an indication of whether a cell or particle hasa predetermined level of reporter and should be chosen for collection.Voltages at the electrodes of the chip can be adjusted or switchedaccording to this determination, for example by signals initiated by orunder the control of a personal computer acting in concert with the LabPC1200 card.

6.12.5. Absorbance Detection

In another embodiment for detecting cells or molecules, absorbancedetection is employed, which typically uses relatively longerwavelengths of light, e.g., ultraviolet (UV). Thus, for example, a UVlight source can be employed. Additional objective lenses can be used toimage a detection region, such that the lenses are preferably positionedfrom the top surface if the PDMS device is made reasonably thin.Measurement of the light transmitted, i.e., not absorbed by the particleor cell, using an adjustable slit, e.g., a Zeiss adjustable slit asdescribed above, is similar to that used in fluorescence detection. Aspectrophotometer may also be used. As molecules, particles or cellspass through the detection window they attenuate the light; permittingdetection of a desired characteristic or the lack thereof. This canimprove the accuracy of the particle sorting, for example, when sortingbased on an amount of a characteristic, rather than presence of thecharacteristic alone, or to confirm a signal.

It is noted that in some cases, detection by absorbance may bedetrimental at certain wavelengths of light to some biological material,e.g., E. coli cells at shorter (UV) wavelengths. Therefore, biologicalmaterial to be sorted in this manner should first be tested first undervarious wavelengths of light using routine methods in the art.Preferably, a longer wavelength can be selected which does not damagethe biological material of interest, but is sufficiently absorbed fordetection.

6.12.6. Optical Trapping

In another embodiment, an optical trap, or laser tweezers, may be usedto sort or direct molecules or cells in a PDMS device of the invention.One exemplary method to accomplish this is to prepare an optical trap,methods for which are well known in the art, that is focused at the “T”intersection proximate to, and preferably downstream of, the detectionregion. Different pressure gradients are established in each branch. Alarger gradient at one branch channel creates a dominant flow ofmolecules, particles or cells, which should be directed into the wastechannel. A second, smaller gradient at another branch channel should beestablished to create a slower flow of fluid from the “T” intersectionto another channel for collection. The optical trap remains in an “off”mode until a desired particle is detected at the detection region. Afterdetection of a desired characteristic, the particle or cell is“trapped”, and thereby directed or moved into the predetermined branchchannel for collection. The molecule or cell is released after it iscommitted to the collection channel by turning off the trap laser. Themovement of the cell or molecule is further controlled by the flow intothe collection well. The optical trap retains its focus on the “T”intersection until the detection region detects the next molecule, cellor particle.

Flow control by optical trapping permits similar flexibility in bufferselection as a pressure driven system. In addition, the pressuregradients can be easily established by adjusting the volume of liquidadded to the wells. However, it is noted that the flow rate will not beas fast when the pressure in one channel is above ambient pressure andpressure in another is below.

6.12.7. Forward Sorting

In an electrode-driven embodiment, prior to loading the wells withsample and buffer and placing the electrodes, the electrode voltages areset to zero. Once the sample is loaded and the electrodes placed,voltages for the driven electrodes are set, for example using computercontrol with software that prompts for the desired voltages, for examplethe voltage differential between the sample and waste electrodes. If thethree wells are equidistant from the “T” intersection, one voltage willbe slightly more than half the other. In a typical run, the voltages areset by the program to start with directing the molecules, particles orcells to the waste channel. The user is prompted for the thresholdvoltage of the PMT signal, to identify a molecule, particle or cell forsorting, i.e. diversion to the collection channel and well. A delay timeis also set. If the PMT voltage exceeds the set threshold, the drivenelectrode voltages are swapped and then, after the specified delay time,the voltages are swapped back. The delay allows the selected molecule,particle or cell enough time to travel down the collection channel sothat it will not be redirected or lost when the voltages are switchedback. As described above, a slight gradient is maintained in the wastechannel, when the voltages are switched, to provide continuity in theflow. This is not strong enough to keep the molecule, particle or cellmoving into the other or “off” channel it if is too close to or is stillat the branch point.

The value of this delay depends primarily on the velocity of themolecules, particles or cells, which is usually linearly dependent onthe voltage gradients. It is believed that momentum effects do notinfluence the delay time or the sorting process. The molecules,particles or cells change direction almost instantaneously with changesin the direction of the voltage gradients. Unexpectedly, experimentshave so far failed to vary the voltages faster than the particles orcells can respond. Similarly, experiments have so far shown that thedimensions of the channels do not effect the delay, on the distance andtime scales described, and using the described electronics. In additionthe speed with which the cells change direction even at high voltagegradients is significantly less than needed to move them down theappropriate channel, when using a forward sorting algorithm.

Once the voltage and delay value are entered the program, it enters asorting loop, in which the ADC of the Lab PC 1200 card is polled untilthe threshold value is exceeded. During that time, the flow of particlesor cells is directed into one of the channels, typically a wastechannel. Once the threshold is detected, the above voltage switchingsequence is initiated. This directs a selected cell or particle (usuallyand most preferably one at a time) into the other channel, typically acollection channel. It will be appreciated that the cells or particlesare being sorted and separated according to the threshold criteria,without regard for which channel or well is considered “waste” or“collection”. Thus, molecules or cells can be removed from a sample forfurther use, or they can be discarded as impurities in the sample.

After the switching cycle is complete (i.e. after the delay), theprogram returns to the ADC polling loop. A counter has also beenimplemented in the switching sequence which keeps track of the number oftimes the switching sequence is executed during one run of the program.This should represent the number of molecules, cells or particlesdetected and sorted. However, there is a statistical chance that twomolecules, cells or particles can pass through simultaneously and becounted as one. In this embodiment, the program continues in thispolling loop indefinitely until the user exits the loop, e.g. by typinga key on the computer keyboard. This sets the DACs (and the electrodes)to zero volts, and the sorting process stops.

6.12.8. Reverse Sorting

The reverse sorting program is similar to the forward sorting program,and provides additional flexibility and an error correction resource. Inthe event of a significant delay in changing the direction of flow inresponse to a signal to divert a selected molecules, cell or particle,for example due to momentum effects, reversible sorting can change theoverall direction of flow to recover and redirect a molecule, cell orparticle that is initially diverted into the wrong channel. Experimentsusing the described electrode array show a delay problem and an errorrate that are low enough (i.e. virtually non-existent), so thatreversible sorting does not appear necessary. The algorithm and methodmay be beneficial, however, for other embodiments such as those usingpressure driven flow, which though benefiting from an avoidance ofelectrical polarities and high voltages, may be more susceptible tomomentum effects.

If a molecule or cell is detected for separation from the flow, andswitching is not fast enough, the molecule or cell will end up goingdown the waste channel with all of the other undistinguished cells.However, if the flow is stopped as soon as possible after detection, themolecule or cell will not go too far. A lower driving force can then beused to slowly drive the particle in the reverse direction back into thedetection window. Once detected for a second time, the flow can bechanged again, this time directing the molecule or cell to thecollection channel. Having captured the desired molecule or cell, thehigher speed flow can be resumed until the next cell is detected forsorting. This is achieved by altering the voltages at the electrodes, oraltering the analogous pressure gradient, according to the principlesdescribed above.

To move molecules or cells at higher velocities, for faster and moreefficient sorting, higher voltages may be needed, which could bedamaging to molecules or cells, and can be fatal to living cells.Preliminary experiments indicate that there may be a limit to thetrade-off of voltage and speed in an electrode driven system.Consequently, a pressure driven flow may be advantageous for certainembodiments and applications of the invention. Reversible sorting may beadvantageous or preferred in a pressure driven system, as hydraulic flowswitching may not be done as rapidly as voltage switching. However, if amain or waste flow can move fast enough, there may be a net gain inspeed or efficiency over voltage switching even though the flow istemporarily reversed and slowed to provide accurate sorting. Pressuredriven applications may also offer wider flexibility in the use ofbuffers or carriers for sample flow, for example because a response toelectrodes is not needed.

6.13. Design and Microfabrication of a Multiparameter Chip

This device incorporates a built-in microfluidic system constructed bymulti-layer soft lithography (27) using techniques such as thosedescribed in Example 9. The integrated fluidic system has an input well,a main channel which incorporates a central hybridization loop, anoutput well, inlet and outlet on/off valves, and a peristaltic pump.Micrographs of an exemplary device are shown in FIGS. 13A and B.

6.13.1. Chip Architecture

In a departure from conventional systems, this embodiment of theinvention does not rely on DNA probes on a substrate that are passivelyexposed to a sample. This chip incorporates a built-in fluidic systemthat actively contacts probes and sample. The fluidic system is made bymulti-layer soft lithography as described herein. See also, Example 9,Unger et al. (6), and U.S. Pat. No. 5,661,222 (32). In this example, GESilicone RTV 615A and 615B are mixed and then poured onto two differentmolds, a fluid or treatment channel mold and an air or control channelmold. Part A contains vinyl groups and catalyst; part B contains siliconhydride (Si—H) groups. The conventional ratio for RTV 615 is 10 parts Ato one part B (10:1). For bonding, one layer may be made with 30A:1B(excess vinyl groups) and the other with 3A:1B (excess Si—H groups).Each layer is cured separately. When the two layers are brought intocontact and cured at elevated temperature, they bond irreversiblyforming a monolithic elastomeric substrate. On these two molds, thereare intrusive (negative) patterns, which define the final indent fluidor air channels in the cured RTV silicone devices. After partial curingin an oven, RTV from the air channel mold is peeled off and placed ontop of the fluid channel mold. With a second baking, these two RTV facesbond together, forming an integrated fluidic system. In this example,the system has two layers, although multiple layers are possible. Airand fluid channels in their respective layers are embedded inside thewhole assembly. Air channels are above and proximate to fluid channelsover some portion of their length. The air channels do not connect withthe fluid channels directly. However, they do interact with each otherat intersections where a microvalve is formed.

In this example the air channels are about 100 or 200 μm wide and about10 μm deep (FIGS. 13A and 13B). Suitable air channel dimensions includethose ranging from about 50-200 μm wide and are from about 2-50 μm deep,preferably about 10-50 μm. A preferred depth is about 10 μm. The fluidchannels are about 100 μm wide and about 10 μm deep. Suitable fluidchannel dimensions include those ranging from about 10-200 μm wide andfrom about 2-50 μm deep, more preferably about 2-20 μm deep. A preferreddepth is about 10 μm.

When sufficient air pressure is applied in the air channel, the RTVmembrane between the air and fluid channels is pushed down, acting as amicrovalve, and restricts or closes the bottom fluid channel. Thepumping pressure or force is not critical. Exemplary pressures are inthe range of 5-50 psi, but any pressure may be used that is sufficientto close microvalves on the chip without damaging the device (e.g.challenging the seal between an elastomeric layer and an adjacent layeror substrate. Fluid speed tends to be determined less by pressure thanby the frequency of valve cycling. Frequencies of about 75-100 Hz,preferably 75 Hz are suitable in certain preferred embodiments. In somepreferred embodiments, valves are restricted but not fully closed,particularly valves comprising a pump. This approach avoids contact byvalve membranes with the opposing face of the channel, e.g. with theprobe substrate. Possible disturbance of the probes is prevented, whilestill providing rapid and efficient pumping action within the loop. Bycontrolling external 3-way pneumatic valves such as LHDA1211111H fromthe Lee Company, (Westbrook, Conn.) the on/off state of each individualmicrovalve on the chip can be manipulated. Three microvalves in a seriesbecome a peristaltic pump when an appropriate on/off pumping sequence isapplied.

The integrated fluidic system has an input well, an output well, inletand outlet on/off valves, a central hybridization or target loop, and aperistaltic pump built in with the central loop. Photographs ofcomponents of an assembled device are shown in FIGS. 13A and 13B.

FIG. 13A shows an input or handling assembly 1 of an integratedmicrofluidic chip. The device has a fluid layer (bottom layer) with aninput mixing T-channel 5, junction 9 and feed channel 15. The airchannel layer (top layer in this example) has six air channels 7,forming a microvalve 17 where each air channel intersects input channel5. The bottom wide 100-μm air channel 3 is used to close the inlet 15via a microvalve created at intersection 13, when the peristaltic pumpat the loop or ring of FIG. 13B starts operating for hybridization. Thefeed channel 15 at the bottom of FIG. 13A connects to the top channel 15of FIG. 13B.

FIG. 13B shows a treatment assembly 20. A treatment layer with fluidchannels is shown. The treatment layer of this example has a target orhybridization loop 28, which in this example forms a circular shape in aplane of the chip. In one embodiment of the invention, the center ringor target loop 28 is used for DNA hybridization with a sample introducedto loop 28 from the feed channel 15. Sample can exit the loop at channel30. This channel can also be closed, for example to selectively keepfluid in a closed loop 28, by using another microvalve (not shown) incooperation with valve 13. Air channels 22 appear as radial fingers inthis example, and can form microvalves 32 where each air channelintersects loop 28. Any three of these channels 22 and valves 32 can beused in series to provide a peristaltic pump.

In this embodiment the control layer and air channels are above or ontop of the treatment layer and fluid channels. This is depicted in FIGS.13A and 13B. The air channels can be seen overlapping the fluid channelsin the view shown in the figures. It should be noted, however, that“up,” “down”, “top” and “bottom” are convenient relative terms. The chipand assemblies of the invention, and the layers and channels, can beoriented in any desired direction. As one example, the device may beflipping over to change a “top” layer or face in a “bottom” layer orface.

Except for air channel 3, which is 100 μm wide, the channel width ofboth the fluid channels and air channels in this example is about 50 μmwide and 10 μm deep. The entire device in this example is about 1″ by 1″in size.

The central loop or detection region 28 of FIG. 13A is where DNAhybridization probes are laid down along the ring, preferably on a glasssubstrate (not shown), following all or part of a path corresponding toall or part of the loop. DNA samples and fluorescent intercalating dyescan enter from the branches of the channel shown in FIG. 13A. On/offstates of each microvalves are controlled by external pneumatic valves(e.g. Lee LHDA1211111H), which either apply 50 or 100-kPa air pressureto the microvalves or vent them to the atmosphere. A cycling frequencyof up to 75 Hz has been demonstrated with complete opening and closingof the valves (75). An alignment structure 24 can be provided on eachlayer of the microfabricated device, to assist in properly aligningadjacent layers when they are overlayed and bonded together. Suchstructures can function in two dimensions, e.g. length×width for visualalignment, or in three-dimensions, e.g. length×width×depth for aphysical or “lock-in-key” alignment.

In this example, channels molded into the RTV are at different depths orlayers, but are exposed to a common face of the chip when peeled offfrom the supporting mold, because the elastomer is transparent. Themultilayer assembly (here a two-layer RTV assembly) is aligned andbonded to a transparent (e.g. glass) substrate (not shown). The bond inthis example forms a hermetic seal when the RTV and glass substrate arecontacted shortly after removing the RTV from the mold. In this example,the glass substrate is patterned in advance with a desired set of DNAprobes. For example, the entire DNA probes for a number of differentdiseases are laid down along a path on the glass substrate thatcorresponds and communicates with the loop 28.

A chip of the kind depicted in FIG. 13 is shown schematically in FIG.17. Fluid lines 101 carry a sample, typically an aqueous solution, intoa detection loop 105. The detection loop 105 is provided with diagnosticspots 110, which for example are DNA or antibody probes affixed to asubstrate and presented to the fluid in the loop channel 105. Controllines 115 are above or below the fluid lines 101, and typically carry agas, preferably air, under pressure. As described herein, microvalve 120is formed where the control lines 115 and fluid lines 101 or 105intersect. Three valves 120 in series along the loop channel 105 providea peristaltic pump in response to an appropriate on/off or open/closesequence. The peristaltic pump circulates fluid within the loop 105. Theloop 105 can be closed or isolated from other fluid channels 101 byclosing the valves on the in and/or out sides of the loop 105.

6.13.2. Chip Fabrication

Air and fluid mother molds were fabricated on silicon wafers byphotolithography. Photoresist (Shipley SJR5740) was spun onto a siliconsubstrate at spin rates corresponding to the desired channel heights.After photolithography, intrusive channels made of photoresist wereformed. Fluid channel molds were baked on a hot plate of 200° C. for 30minutes so that the photoresist could reflow and form a rounded shape,which in this embodiment is important for complete valve closure (76). Aone minute trimethylchlorosilane (TMCS) vapor treatment was applied tothese molds before each RTV replication process to prevent adhesion ofcured RTV to the photoresist. With this protective coating, molds can bereused many times. A mixture of GE RTV 615 components, in an RTV ratioof 615A:615B of 30:1 was spun on a fluid channel mold at 2,000 RPM,which covers the photoresist channel and leaves a thin membrane on topof it. At the same time, 3:1 GE-RTV 615A:615B was poured onto an airchannel mold. After baking both molds in an oven at 80° C. for 20minutes, the block of 3:1 RTV with air channels at the bottom was peeledoff from the second mold. Air supply through-holes were punched. Alignedto the fluid pattern under a microscope, the air channel layer was thenpressed against the thin 30:1 RTV on the first mold. A post-bake of anhour at 80° C. made the two silicone pieces or layers chemically bond toeach other. After peeling the assembly off from the remaining mold andpunching the fluid through-holes, the monolithic RTV device could sealhermetically to a glass cover slip. This glass cover slip can bechemically patterned in advance, to make an active diagnostic chip. Ifhigh-resolution transparency photomask are used (minimum feature size:˜10 μm), the whole process from the design to the final products can beaccomplished very quickly, even without automation, e.g. in one day.

6.13.3. Operation of a DNA Hybridization Embodiment

In operation, a few microliters of sample, containing perhaps as few as50 or 100 molecules of DNA, are loaded from an input well (not shown)via inlet 5 of handling assembly 1, and fill the fluid channels of thedevice by capillary action. After the central loop 28 is completelyfilled, inlet and outlet valves (not shown) associated with channels 26and 32 are closed. The peristaltic pump is turned on, via microvalves22, to move the fluid around and around loop 28 in a closed circle.Instead of a passive diffusion process, the target DNA fragments,polynucleotides or molecules contained inside the sample are activelypumped to pass each individual hybridization spot. The sample passedevery probe several times, so that almost all DNA that targets a probewill find the right spots to hybridize with. Very little sample isneeded. This is a significant improvement over conventional passive DNAchips.

Improved Accuracy. Intermittent heating can also be applied to denaturefalse hybridization and thus obtain an even more accurate diagnosticresult. This approach improves the signal-to-noise ratio in certainembodiments. Another technique to improve accuracy, and (for example) toavoid false positives, is to provide additional hybridization spots,before or within the target loop, to extract common DNA, leavingunmatched sample DNA to bind with target probes on other hybridizationspots in the detection loop, for labeling and/or detection. In this way,DNA that is known not to match any of the target DNA probes can bescreened or filtered out.

Optical Detection. After hybridization, the chip can be checked under anoptical microscope easily, because the whole body of the chip istransparent. Intercalating dyes, incorporation of fluorescent-labeledsingle nucleotides, DNA beacons or other well-established detectionschemes can be used to determine the final diagnostic results.

In one preferred embodiment, intercalating dye is used. Fluorescentintercalating dyes, such as YOYO-1, TOTO-1 and PicoGreen from MolecularProbes, have been demonstrated to have, very high affinity todouble-stranded DNA (dsDNA), big excitation cross-sections and highquantum efficiency. Most important of all is that their fluorescence isenhanced more than one-thousand fold when bound to dsDNA fragments (7,8), and shows a relatively high selectivity to dsDNA compared tosingle-stranded DNA probe (ssDNA). When concentration of dsDNA is below100 pg/ml, 10× more concentration of ssDNA results in no more than a 10%change in the signal intensity of PicoGreen stained DNA (10). This meansthat at least 100-fold discrimination between dsDNA and ssDNA has beendemonstrated. With such a high specificity and binding enhancement,these dyes are highly suitable as an indicator of hybridized DNA probes.Detection of individual stained dsDNA molecules using laser excitationhas been reported, e.g. (8, 9), but is not necessary. An opticalmicroscope with a mercury illumination lamp and a CCD camera gives areasonable justification between hybridized and non-hybridized spots aslong as a few seconds of exposure is applied to obtain a similar signallevel.

In one embodiment, the target loop (e.g. loop 28) is provided with asecond inlet valve and/or inlet channel, to deliver an additional flowor additional materials or reagents, such as a buffer or sample-freemedium. After hybridization, this valve (not shown) and an outlet valvefor channel 30 (not shown) are opened. Buffer with dye molecules canthen be introduced to flush the loop 28, to remove free DNA in thesolution, and also to stain the hybridized fragments at the bottom.After a few minutes of incubation, the whole chip can be checked underan optical microscope, as described herein. A computer program can beused to determine the existence of certain disease targets by athreshold algorithm. Severity of the infection can be also determinedaccording to the fluorescent intensity of the correspondinghybridization spot.

DNA beacons are a special kind of hybridization probes, and arecommercially available for example from Integrated DNA Technologies,Inc, (Coralville, Iowa); Oswel Rearch Products Ltd. (UK), Cruachem, Inc.and Research Genetics, Inc. When not hybridized with their complimentaryDNA fragments (target DNA), they are self-annealed to themselves andthus quench their own fluorescence. Therefore, no additional stainingstep is required for a final diagnosis. However, these probes are moreexpensive, and special probing sequences have to be carefully chosen inorder for the self-annealing mechanism to take place. In-situ enzymaticlabeling with fluorescent molecules is another well-known method and isobviously compatible with the design of this lab-on-a-chip device. Theseand other reporter and labeling techniques can be used in concert withthe microfluidic devices and methods of the invention.

In addition to the diagnostic design described above, other functionscan be incorporated into the integrated device as needed. Switchingvalves and mixing chambers can be designed and built into the chip. Forexample, an automatic inline restriction and denaturing process can beincluded before the hybridization process, to reduce handling labors.For an extremely small quantity of target samples, a PCR chamber canbuilt into the device, and reactions can be carried out by an externalor an inline thermal cycler (11). Other enzymatic labeling and reactionscan be easily incorporated into it just like the way we described abovefor DNA fluorescent staining. A device of the invention is shownschematically in FIG. 14. A microfluidic assembly 40 comprises a fluidor treatment layer 44 and a control or air layer 46. In this embodiment,fluid channels are microfabricated into the treatment layer 44, and airchannels are microfabricated into the control layer 46. The controllayer is on top of and is bonded or sandwiched to the treatment layer.These layers overlap, and are typically but not necessarily coplanar orcoextensive. In this embodiment the layers are transparent, and thecontrol layer is bonded on its upper face, in whole or in part, to atransparent cover layer 48, preferably glass (e.g. pyrex).

In the treatment layer 44, a sample inlet channel 50 communicates with atarget hybridization loop 58, which in turn communicated with an outletchannel 68. A reagent channel 52 communicates with the sample inletchannel 50. The control layer 46 has a first air channel 54, which formsa microvalve 72 to open and close the sample inlet channel 50,particularly with respect to the target loop 58. Similarly, a second airchannel 66 provides a microvalve 74 to open and close the outlet channel68. At least three pump channels 56 providing microvalves 78 are on thecontrol layer. Using air in the pump channels 56, e.g. by changing thepressure, the valves 78 can be opened and closed in a series or cycle,to create a pumping action in the target loop 58, represented by thefour arrows in FIG. 14 showing a counterclockwise motion. The heavyarrows represent fluid, e.g. sample, in the fluid channels. The lighterarrows represent air in the air channels, e.g. to actuate themicrovalves.

The cover layer 48 (e.g. a glass substrate) is patterned with targetmolecules (e.g. DNA probes) in distinct spots 60, 64 on the inward faceof the cover layer, i.e. toward the other layers. The pattern of spotsfollows a path that corresponds to the path of the target loop 58, andthe fluid channel comprising the loop 58 is open or exposed to, andtypically is sealed by, the inward face of the cover layer. Thus, thecover layer 48 forms a ceiling, wall or floor of the fluid channels,depending on the orientation Of the assembly 40. In the view shown inFIG. 14, the cover layer 48 forms a ceiling for loop 58, and may beextended to cover other channels, the entire assembly, or may extendbeyond the assembly (not shown). In a preferred embodiment, all threelayers (treatment, control and cover) are coplanar and coextensive).

In operation, a sample is loaded into inlet channel 50, and the fluidchannels are allowed to fill by capillary action. Channel 54 may be usedto add reagents or buffer, or to wash the fluid channel, as and whendesired. The microvalves 72 and 74 are, closed, to isolate the loop 58.At least three microvalves 56 are actuated to form a peristaltic pump,which circulates sample inside the loop, for repeated and activeexposure to the probes at the hybridization spots 60, 64. The blackspots 60 represent a component of the sample (e.g. a DNA fragment)binding to a corresponding target probe on the cover layer. The whitespots 64 represent target probes fixed to the cover layer and exposed tocirculating sample.

In a preferred DNA hybridization embodiment, the chip is washed (e.g.with buffer), loaded with sample, the loop is closed, the sample iscirculated around the loop, and the chip is heated up to a temperaturethat is near but below the lowest annealing temperature of the probeDNA. The heat acts to prevent or reverse false hybridization, bybreaking weak bonds between sample molecules and non-matching probes,without denaturing the DNA of the sample or the probes. Each primer hasan annealing temperature, which can be determined or calculated by knownmeans, for example using commercially available probe design software.Mismatched hybridization may be a function of annealing temperature, andaccordingly, maintaining a temperature that is below the annealingtemperature tends to minimize false hybridization. Typically, theannealing temperatures are such that the operating temperature is set inthe range of about 55-70° C. The hybridization cycle is completed byopening valve 74 to drain the loop of sample, following by a washingstep, in which buffer or other reagent in fed to the loop from channel52 by opening valve 72. Washing conditions can be optimized to removesample, leaving behind the positive hybridization spots 60, where samplemolecules are bound to or associated with target probes. Alternatively,washing conditions can be optimized to flush the loop for a fresh roundof sample and testing. Loading, treatment, detection, and washing cyclescan be repeated as desired, in any order, for any length of time, andaccording to any protocol.

In one embodiment, fluorescent labeled nucleotides and polymerase areintroduced to the loop (e.g. via channel 52 and valve 72) to extend thehybridized samples and provide a detectable reporter, followed by awashing step. For example, a positive hybridization can be detected byobserving any hybridization spots that fluoresce, for example using anoptical microscope. The microscope can be used with a charge coupleddevice (CCD) to image the loop, and processing, imaging, and analysiscan be automated or assisted by a computer, e.g. a personal computer.

This method is suitable for samples of very low volume, e.g. 1-50 μl,typically 10 μl. In certain embodiments it may be desirable to dilute orconcentrate the sample. For example, in embodiments where primerextension is done on the chip (at the hybridization spots), dilution oradditional samples may be needed If the extension length of a primer isabout 5K, hybridization can be imaged (e.g. using a CCD) from as few asabout 10 extended DNA molecules on each spot. Accounting for possibleloss of sample, e.g. during device or target loop loading, a minimum ofabout 20 molecules in the loaded sample is preferred. More typically, aminimum of about 30 to 100 molecules is preferred.

6.14. Surface Patterning of a Multiparameter Chip

This example describes the surface patterning of a microfluidic deviceof the invention, using two different kinds of surface chemistries.These devices are useful, for example, to measure gene expression anddetect the presence of pathogenic DNA. As described above, themicrofluidic devices pump solutions of target DNA over a set of anchoredprobes in order to ensure that all of the target DNA is exposed to eachof the probes. This provides increased sensitivity as well as decreasingthe amount of time needed for hybridization. Chips with high sensitivityare useful for measuring single cell gene expression. This highersensitivity may eliminate the need for PCR in many cases, e.g. ofpathogen detection. Such devices can provide multiple analysis ordisease diagnosis with one integrated lab-on-a-chip.

The microfluidic devices of the invention are compatible in structure,material and manufacturing with the delicate surface chemistry requiredto anchor or synthesize DNA probes onto the chip. The desired patterningor surface chemistry is used to place probes on the chip, for example ina detection or target loop as described above. Small amounts of materialare manipulated in order to perform biochemical reactions on the chip,including the ability to pump a sample over probes that are patterned onthe chip.

The soft lithography techniques described herein are suitable forhandling DNA and reactions involving DNA, and for patterning a substratewith DNA probes. The elastomeric devices of the invention provide anumber of important advantages over conventional micromachining, such asease of fabrication, room temperature sealing of devices to glasssubstrates, good optical properties, and low materials cost.Microfluidic networks fabricated in such a manner can easily be sealedto substrates with delicate surface chemistry. Another aspect of softlithography is the ability to chemically pattern surfaces using fluidflow (78, 79). The multilayer fabrication process in silicone elastomer,described herein, furnishes the easy fabrication of devices with movingparts, including microfluidic valves and pumps (U.S. Patent ApplicationSer. No. 60/186,856, filed Mar. 3, 2000 entitled “MicrofabricatedElastomeric Valve and Pump Systems”)

In this example, a microfluidic device of the invention has surfacesthat are chemically patterned with biotin/avidin and DNA using fluidicnetworks that are compatible with further fluidic processing. Pumps canbe incorporated into the device to both meter reagents and pump fluid ina closed loop.

Multilayer soft lithography, described herein, is used to make 3-Dmonolithic elastomer devices with a combination of air and fluidchannels. The devices of this example were made as described in Example13. When an air channel passes above another fluid channel, the thinmembrane between these two channels becomes a valve. By applying airpressure in the air channel, the membrane collapses and stops the fluidflow. Releasing the pressure then re-opens this valve. Three valves inseries become a peristaltic pump when an appropriate on/off airpressures are applied in a sequence. For example, three valves in seriescan be represented by the letters “XYZ,” with 0 representing a closedvalve and 1 representing an open valve As shown in the following table,the XYZ sequence 100, 110, 010, 011, 001, 101 pumps water to the right,e.g. from opened (on) valves toward closed (off) ones. These and othersequences can be used to direct and control fluid flow, change flowdirection, start and stop flow, etc. For example, the table belowspecifies an exemplary sequence of six steps or “words” that can be usedto pump fluid through a microfabricated rotamer loop.

Valve Step X Y Z 1 1—on 0—off 0—off 2 1—on 1—on 0—off 3 0—off 1—on 0—off4 0—off 1—on 1—on 5 0—off 0—off 1—on 6 1—on 0—off 1—onThree word pumping sequences may also be used, for example the sequence001, 010, 100 or, alternatively, the sequence 011, 110, 101. Performancebetween these different sequences is relative similar when theperistaltic pump is operated at the linear regime; i.e. when the cyclingfrequency is less than about 75 Hz.

A schematic diagram of a peristaltic pump of the invention is shown inFIG. 15. In this exemplary embodiment, the distance between the airchannels and the fluid channels, where they intersect, is a vertical gapof about 30 μm. The fluid and air channels are preferably disposed at anangle to one another with a small membrane of elastomeric materialseparating the top of one channel (e.g. an air channel) from the bottomof another channel (e.g. a fluid channel).

Two independent methods of surface patterning are disclosed. The firstmethod provides patterning of the protein streptavidin, a commonbiochemical “glue” that binds biotin with nearly covalent strength.Using the streptavidin surfaces, biotin-labeled reagents are selectivelyanchored, including proteins and nucleic acids. The second methodprovides direct attachment of amine-modified DNA molecules to a surfaceusing a commercially available surface chemistry from the companySurmodics, Eden Prarie, Minn.

6.14.1. Straptavidin Binding

In the first method, half of the surface of a glass cover slip (VWR #1,from VWR Scientific Products, Inc., Chester, Pa.) was derivatized withbiotin. The coverslip was aligned and contacted with the fluid channelsof a silicon elastomer layer, as described above, and the channels wereflowed with avidin-fluorescein conjugate, after which the channels wereflushed with water and removed from contact with the coverslip. Thecoverslip was washed. As shown in FIG. 16A, the avidin molecules boundto the derivatized part of the glass surface with a high affinity in theregions defined by the channels, forming fluorescent detectable stripesin a distinctive line pattern. Regions which were not derivatized withbiotin function as a control and showed a much lower level of avidinbinding.

The substrate surface can be successively patterned. In this embodiment,the surface was patterned with non-fluorescent streptavidin by bondingan elastomeric device with channels to the substrate and flowingstreptavidin down the channels. As before, streptavidin boundselectively to the surfaces that were exposed to the channels, and notto the elastomer surfaces. Since streptavidin is a tetramer, eachmolecule has at least two exposed groups free to bind more biotin. Thiswas demonstrated by removing the elastomeric channels and re-bonding inan orientation that was rotated by 90 degrees. Biotin-fluoresceinconjugates were flowed down the channels and then washed with water. Asshown in FIG. 16B, the fluorescent biotin binds selectively to theregions that are derivatized with streptavidin. A checkerboard patternwas obtained in this figure by flowing streptavidin horizontally (200μm) and biotin-fluorescein conjugates vertically (100 μm).

6.14.2. Covalent Immobilization of DNA

Surfaces can be prepared and patterned with DNA using commerciallyavailable silanized slides, such 3D-Link, provided by Surmodics. In thisexample, DNA samples were prepared by PCR of a 2 kpb region of lambdaphage DNA using amino-terminated primers. The DNA was attached in situby flowing it through an elastomeric channel replica made from the airchannel mold of the diagnosis chip, i.e. the finger patterns in FIG.13B. After overnight incubation, the elastomeric device was peeled offfrom the slide, and washing and immobilization steps were followedaccording to the manufacturer's protocol. To show that the DNA wasattached and patterned to the surface, a diagnostic RTV device (as shownin FIGS. 13A and 13B) was aligned and attached to the same slide. Then,the DNA intercalating dye PicoGreen (Molecular Probes P-7581) was flowedthrough the bottom fluid channels 15, 28, 30, of which the central ring28 intersected with every DNA finger pattern on the slide. Theintersection of the channels fluoresced, as shown in FIG. 16C. Althoughthis particular example uses uses DNA, protein-binding assays and othermolecular affinity assays can also be used with these fluidic systems.

The DNA diagnostic chip in this example has a junction 9 for mixing andmetering reagents (FIG. 13A), which then leads into a fluidic loop 28.In this embodiment, probe molecules are anchored or immobilized in theloop, via bonding to an aligned substrate, so that the sample (or probetarget) can circulate around. The loop has air channels 22 formingperistaltic pumps to control circulation. The fluidic connections intoand out of the loop 28 are controlled by input and output valves,respectively. (In other embodiments, probes may circulate freely in atarget loop, or may be fixed to a surface such as beads, for circulationin the target loop, exposure to sample, and later imaging or detectionat a detection region of the device.)

Mixing and metering occurs in the first part of the chip (FIG. 13A). Asolution containing fluorescent dye is mixed with an aqueous solution,for example using flows that can be alternated with different (e.g. two)valve-firing schemes. If the valves are opened and closed in synchrony,the fluid mixing is controlled by diffusion. There are two segregatedflows when the valves are open, which quickly mix by diffusion when thevalves are closed. If the valves are opened and closed alternately,slugs or droplets of fluid are injected into the stream.

Fluid is circulated within the closed loop using the peristaltic pumpson the perimeter. The channels were loaded with fluorescent beads (2.5μm in diameter). The beads could be visualized as the fluid circulatedaround the loop and clearly showed rotary motion with no net flux intoor out of the loop. Thus reagents can be repetitively exposed todiagnostic probes anchored on the surface, and their binding is notlimited by diffusion. All or substantially all target DNA in a sample iseventually captured by corresponding probes after several passages. Sucha device can also be used to rapidly mix viscous liquids, since theparabolic flow profile of the fluid will tend to “wrap” the two fluidsaround each other.

In these experiments, surfaces were chemically patterned withbiotin/avidin and DNA using fluidic networks in a way that is compatiblewith further fluidic processing. Experiments also show how pumps can beincorporated into an integrated device to both meter reagents and pumpfluid in a closed loop. Fast in-line mixing and rotary pumping weredemonstrated to overcome the slow hybridization process.

In a microfluidic device with a closed loop and peristaltic pump, asdescribed, avidin coated 1 μm beads were captured onto biotin spotswithin 4 minutes after the pump was activated. In contrast, when usingpassive diffusion, only beads very local to the reaction spots werecaptured even after hours of waiting time.

Rotary flow in a closed loop cannot be achieved by the electrophoresisor electroosmotic flow used in the conventional and most commonlab-on-a-chip devices, because of the existence of two electricpolarities. The problem of buffer depletion due to electrolysis inelectroosmotic or electrophoretic flow control does not exist in pumpand valve devices. However, it should be noted that electroosmotic orelectrophoretic flow control systems, and/or other flow control systems,can be used in combination with the pumps and valves disclosed herein.For example, other flow control systems can be used to move fluids onother parts of a multifunction chip, or from one chip to another in adevice comprising cooperating microfluidic chips, layers, units orsubunits.

6.15. In-Line Rotary Mixing

The exemplary device of FIG. 13 has a junction 9 for mixing and meteringreagents, which then leads into a fluidic loop 28 via inlet channel 15.In this example, mixing and metering in the first part of the chip (FIG.13A) is described. A solution containing fluorescent dye was mixed ormetered with an aqueous solution. For example, the flows can bealternated with two different valve-firing schemes. If valves 7 areopened and closed in synchrony, fluid mixing is controlled by diffusion.Two segregated flows are observed when the valves are open. These flowsquickly mix by diffusion when the valves are closed. If the valves areopened and closed alternately, then slugs or droplets of fluid areinjected into the stream.

Fluid was pumped within the closed loop 28 using the peristaltic pumpson the perimeter. The channels were loaded with fluorescent beads (2.5μm in diameter). The beads could be visualized as the fluid circulatedaround the loop and showed rotary motion with no net flux into or out ofthe loop. Thus reagents can be repetitively exposed to diagnostic probesanchored on the surface, and their binding will not be limited bydiffusion. Likewise, all target molecules in a sample (e.g. DNA) areexposed to and eventually (and relatively quickly) interact with or arecaptured by their corresponding probes after several passages or turnsaround the loop. There is little or no loss of sample, and in many oreven most cases PCR amplification of the sample is not needed.

Mixing within the loop can also be done without the presence of probesin the loop, for example to facilitate a chemical reaction orcombination of different flows or ingredients in different flowsintroduced upon a mixing protocol as described above.

The loop can contain or be provided with any material or reactant,immobilized or not, for any reaction or interaction with any othermaterial or reactant provided to the loop. Immobilized reactants may beattached to any substrate, fixed or mobile, including carrier molecules,beads, or a substrate (e.g. glass) communicating with the elastomericloop channel 28.

A device of FIG. 13 was also used to rapidly mix viscous liquids. Theperistaltic pump provides a flow profile (e.g. parabolic) whereby two(or more) fluids tend to “wrap” around each other to provide in-linerotary mixing, as shown in FIG. 18.

No Pumping Action. In FIG. 18A, buffer containing fluorescent beads camein from the left input channel at the T-junction while buffer containingthe fluorescent dye FITC came in from the other side (FIG. 18A). Becauseof a laminar flow profile, these two fluids did not mix with each otherand actually split the flow channel into halves, one side with onlybeads and the other side with fluorescent dye. When they enter thecentral ring or loop, without pumping, the ring was also split into twodistinct parts. On the left-hand side, there were just beads flowingthrough and on the right-hand side, there was only bright and uniformfluorescent dye. At the bottom of the ring, these two flows met witheach other again (FIG. 18A). The channel was split into two distinctportions again. The inset illustrates the flow pattern shown in thephotograph.

Pumping Activated When peristaltic pumping was turned on at the centralring, the situation changed significantly. As shown in FIG. 18B, bothdye and fluorescent beads were well mixed at the output channel. Part ofthe fluid was actually pumped back to the input of the ring and forcedthe two distinct streams to mix with each other. This fast in-linemixing by rotary pumping is useful in many microfluidic systems,particularly where time and space is critical, and when fluid containssubstances with small diffusion constants, such as DNA and micron-sizedbeads.

6.16. Model Biotin and Avidin System

A biotin/avidin model system demonstrated the difference, in terms ofdetection efficiency, between a passive and an active diagnosis chip. AnRCA-cleaned cover slip was first patterned by flowing biotinylationsolution though an attached RTV device with eight finger channels madefrom an air-channel mold as in FIG. 13. After overnight incubation, theRTV device was peeled off and the cover slip was washed with DI water. Amultiple disease diagnosis device (FIG. 13) was then attached with itscenter aligned to the center of the biotinylated fingers. Therefore, thecentral ring was able to intersect with all biotin fingers and formedeight diagnostic spots. The pattern shown in FIG. 16C was made in thesame manner except that the DNA molecules were anchored on the surfaceinstead of biotin molecules.

After this biotin/avidin diagnosis device was made, 1-μm fluorescentbeads (F-8776 from Molecular Probes) coated with NeutrAvidin, aderivative of avidin with less nonspecific binding, were introduced intothe mixing loop or ring from the input channel. Once the ring wasfilled, the flow was shut off right away. Because of the strong affinitybetween avidin and biotin, beads that were close to the biotinylatedspots are “grabbed” onto them and show positive diagnostic signals.

In trials without circulation in the loop (without rotary pumping), andthus under the action of passive diffusion only, no difference of beadconcentration between biotinylated spots and the rest of the channel wasobserved even after we waited for 30 minutes. In one experiment, thefirst appearance of differentiated biotinylated spots was not observeduntil after four hours. Most of the beads reaching within a distance ofabout 50 μm to the biotin pad or spot were grabbed onto it. This is avery slow process, and beads on one side of the ring would have littleor no chance to get onto the (biotinylated) detection spots on the otherside, a necessary condition for sensitive multiple disease diagnostics.(The diffusion constant D of 1-μm beads is ˜2.5×10⁻⁹ cm²/s (85), whichis 40 times slower than 1-kpb DNA molecules (D is ˜1×10⁺⁷ cm²/s). So, 4hours for the beads would be about 6 minutes for 1-kbp DNAmolecules—still a slow process for diffusion across a 50 μm space. Thisparticularly so in comparison with the active pumping scheme, which cancover several millimeters in 4 minutes. At least two orders of magnitudein speed can be achieved even for 1-kbp DNA molecules.)

When the peristaltic pump was activated, beads were actively moved inthe loop now. Within 4 minutes, more than 80% of the beads in thecentral ring were quickly grabbed onto biotinylated spots, as shown inFIG. 19. Thus, the active detection scheme of the invention providessignificant advantages in speed and efficiency. These devices andmethods can also be advantageously used to provide improved (e.g.faster, more accurate and less costly) affinity purification systems.

6.17. Operation of a Multiparameter Chip

A few microliters (e.g. 1-50 μl) of sample are loaded from the inputwell and fill the fluid channels by capillary action. After the centralloop is completely filled, inlet and outlet valves are closed and theperistaltic pump is turned on which move the fluid around in a circle.Instead of passive diffusion process as used in conventional chips, thetarget DNA in the sample is actively pumped to pass each individualcomplementary fragment on the surface of the substrate. With smallchannel dimension, 100 μm×10 μm typically, the hybridization rate andefficiency are enhanced significantly (2). The size of the channels istypically 50 to 100 μm wide and 10 μm deep. It takes only 10 seconds fora 1-kbp DNA fragment to diffuse 10 μm, all the way from the top of thechannel to reach the hybridization probes at the bottom.

Moreover, since the sample will pass every probe several times, almostall target DNA will locate and hybridize with the corresponding DNAprobes. Also, very little sample will be wasted during thishybridization process, which is a significant improvement overconventional passive DNA chips. Hybridization of the sample tocomplementary DNA probes is easily visualized under an opticalmicroscope because the whole body of the device is transparent.Intercalating dyes, incorporation of fluorescent-labeled singlenucleotides, and DNA beacons or other well-established detection schemescan be used to determine the final diagnostic results. Examples ofFluorescent dyes, particularly those that intercalate between thepolynucleotide backbone, include, but are not limited to, Hoechst 33258,Hoechst 33342, DAPI (4′,6-diamidino-2-phenylindole HC1), propidiumiodide, dihydroethidium, acridine orange, ethidium bromide, ethidiumhomodimers (e.g., EthD-1, EthD-2), acridine-ethidium heterodimer (AEthD)and the thiazole orange derivatives PO-PRO, BOPRO, YO-PRO, TO-PRO, aswell as their dimeric analogs POPO, BOBO, YOYO, and TOTO. The dimericanalogs, especially YOYO-1 and TOTO-1, are particular suitable for usewith the present invention due to their high binding affinity fornucleic acids, which results in extremely high detection sensitivity.All of these compounds can be obtained from Molecular Probes (Eugene,Oreg.). Extensive information on their spectral properties, use, and thelike is provided in Haugland, 1992, incorporated herein by reference.

Fluorescent intercalating dyes, such as YOYO-1, TOTO-1 and PicoGreen(Molecular Probes) are generally preferred because they have beendemonstrated to have very high affinity to double-stranded DNA (dsDNA),large excitation cross-sections and high quantum efficiency. Theirfluorescence is enhanced more than 1,000 fold when bound todouble-stranded DNA (dsDNA) fragments (7, 41) and they have relativelyhigh selectivity to dsDNA compared to single-stranded DNA probe (ssDNA)(3). This high specificity and binding efficiency makes them verysuitable for use as an indicator of hybridized DNA probes. Detection ofindividually stained dsDNA molecule using laser excitation has beenreported in many places (7, 8). However, an optical microscope with amercury illumination lamp and a good CCD camera gives a reasonablejustification between hybridized and non-hybridized spots as long as afew seconds of exposure is applied to obtain a similar signal level(FIG. 9).

A second inlet valve incorporated within the central hybridization loopis used because, after hybridization, this valve and the outlet valvecan be opened and buffer with dye molecules can be flushed into theloop. This flushing action removes DNA molecules that do not hybridizeto the DNA probes and are therefore free in the solution, but which willalso stain the hybridized fragments retained in the loop. After a fewminutes of incubation, the whole chip can be checked under an opticalmicroscope as described above. A computer program can be used todetermine the existence of certain disease targets by a thresholdalgorithm. Such algorithms are known and can be determined empirically,for example by comparing the fluorescence of a probe and samplecombination (in a fluorescent reporter embodiment) with a knownreference standard. Severity of the infection can also be determinedaccording to the fluorescent intensity of the correspondinghybridization spot.

Commercially available DNA beacons are very useful as hybridizationprobes because if they do not hybridize with their complimentary DNAfragments (target DNA), they self-anneal to themselves and thus quenchtheir own fluorescence. Therefore, no additional staining step isrequired for the final diagnosis. In-situ enzymatic labeling withfluorescent molecules is another well-known method and is obviouslycompatible with this device.

Polypeptides such as antibodies, antigens, receptors etc. can also becoupled to surface of the solid substrate of the chip. Examples offluorescent dyes that can be coupled to these proteins are discussed inExample 7.

6.18. Additional Embodiments 6.18.1. Additional Structures and Functions

The analysis unit of the invention, including a target loop detectionregion, can be combined with other structures and features on one ormore chips, in an integrated design. In addition to diagnostic designsdescribed above, additional functions can be incorporated into theintegrated device as needed. Switching valves and mixing chambers can beeasily designed and built into it. An automatic in-line restrictiondigest and denaturing process can be included before the hybridizationprocess to save the handling labors. For an extremely small quantity oftarget samples, a PCR chamber can also be built into the device and thePCR reaction can be easily carried out by an external or an inlinethermal cycler. The thermal cycler can also be used to applyintermittent heating to reduce non-specific binding during thehybridization process and thus obtain more accurate diagnostic results.Other enzymatic labeling and reactions can be easily incorporated intothe device as described above for DNA fluorescent staining.

6.18.2. Additional Loop Channel Shapes and Geometries

Although in one preferred embodiment of the invention, illustrated inFIG. 13B, the loop channel in a microfluidic device forms a circularloop, a loop channel may actually form any shape of loop. For instance,FIG. 14 shows another exemplary embodiment where the loop channel formsa rectangular (e.g., a square) loop. The invention also providespreferred embodiments, however, wherein the loop channel forms a shapethat optimizes the length of the channel through the loop. Inparticular, the invention provides preferred embodiments where the loopchannel forms a shape that increases or maximizes the perimeter aroundthe loop.

In more detail, in many embodiments of the invention it is desirable toincrease the length of a loop channel in a microfluidic device. Forexample, in embodiments where the loop channel contains one or moretarget molecules (for example, one or more nucleic acid probes, or oneor more antibody probes, etc.), by increasing the length of a loopchannel a user may simultaneously increase the number of targetmolecules within the loop channel, thereby increasing the number ofdifferent molecules (e.g., different nucleic acid sequences or peptidesand/or polypeptides) in a sample that may be detected. Conversely,however, the microfluidic devices of the invention are preferably small(e.g., between 0.5 cm and 5.0 cm on each side, more preferably about 1cm on each side, and between about 0.1 mm and 10 mm thick). The loopchannel in a microfluidic device, therefore, preferably forms a loopenclosing a limited area. For example, in preferred embodiment the loopcovers an area that is no more than about 5 cm long on either side, morepreferably is no longer than about 1 cm on either side, and still morepreferably is less than about 5 mm long on either side. In oneparticular embodiment, for example, a microfluidic device of theinvention is designed for combination with standard 96-well or 384-wellmicrotiter plates. In such embodiments, a loop channel preferablyencloses an area no wider than the area of an individual microtiter well(e.g., 9 mm×9 mm for 96-well plates, or about 4.5 mm×4.5 mm for 384-wellplates). It is therefore extremely desirable to configure the loopchannel in a microfluidic device of this invention so that the channelforms a loop having the greatest possible perimeter within a given area.Such loop geometries allow a user to increase or maximize the length ofthe loop channel while simultaneously confining the loop to the verysmall dimensions of a microfluidic device.

FIG. 20 illustrates one exemplary embodiment of such a preferred loopchannel geometry. The loop channel (201) encompasses a rectangular(e.g., square) area, and comprises a plurality of interconnectedmicrofluidic channels (202-210). Microchannels 202-209 are a series ofparallel and antiparallel channels, connected at their ends, that coverthe rectangular surface area of the loop. Microchannel 210, which isperpendicular to the plurality of parallel and antiparallel channels202-209 runs along the outer parameter on one side of the rectangularloop area.

As used here to describe channel geometries, the terms parallel andanti-parallel refer to channels that do not intersect and through whichfluid travels in opposite directions (or is designed to travel inopposite directions). In more detail, a pair of microchannels is saidhere to be parallel if the two channels do not intersect, and eitherfluid flows through the channels in the same direction or the channelsare designed so that fluid flows through the same channels in the samedirection. A pair of microchannels is said here to be antiparallel ifthe two channels do not intersect, and either fluid flows through thetwo channels in the opposite direction or the channels are designed sothat fluid will flow in opposite directions through the two channels.Thus, assuming a counter-clockwise flow of fluid through loop channel201, fluid travels from left to right through microchannels 202, 204,206, and 208. These channels are therefore said to be parallel to eachother. Fluid travels from right to left through microchannels 203, 205,207 and 209. These odd numbered channels are therefore antiparallel tothe even number channels shown in FIG. 20. It is noted that thedesignation of any pair of channels in FIG. 20 as parallel oranti-parallel will not be affected by the direction of fluid flow (e.g.,either clockwise or counterclockwise) through loop channel 201.

6.18.3. Additional Uses

Besides diagnosis of infectious diseases, such as tuberculosis,hepatitis and HIV, the device can be also used for detection of humangenetic defects, such as cystic fibrosis, phenylketonuria and breastcancer genes (11). Moreover, multiple chemical reactions and otherbiological diagnosis can also be done using an appropriate set of probesand a suitable operating protocol. All of these can be considered asextensions of its applications.

6.18.4. Cell Analysis

In another embodiment, a cell or tissue sample can be processed andanalyzed on an integrated chip. Cells are introduced to a firsttreatment chamber from a well or reservoir, on or off the chip. Thefirst chamber includes a soap or other reagent to break or lyse the cellmembrane. In a second treatment chamber the lysed cell material istreated with a digestion enzyme. In a third chamber, cell debris andprotein can be washed away, leaving denatured (fragmented) DNA, which isthen delivered to a target loop detection region, as described above.Also, magnetic beads coated with mixed desired probes can be used tohybridize and pull down DNA of interest, or alternatively, a pool of DNAthat is not of interest, leaving other DNA for further analysis.

Besides diagnosis of infectious diseases, such as tuberculosis,hepatitis and HIV, the device can be also used for detection of humangenetic defects, such as cystic fibrosis, phenylketonuria and breastcancer genes (12). Multiple chemical reactions and other biologicaldiagnosis can also be done using an appropriate set of probes and asuitable operating protocol.

The lab-on-a-chip device of the invention uses a sample size that isseveral orders of magnitude less than is needed for conventionalmethods. Instead of many cubic centimeters or “ccs” of a blood sample, afew droplets (2-100 μl) is sufficient. The active design of the devicespeeds up the detection process significantly. Multiple diseasediagnostics can be done in just minutes. This device is also economicaland disposable due to the material and the easy molding process.Automatic computer control can be easily integrated by controlling theexternal switching of pneumatic valves via electronic driving circuits.Manual labors and chances of errors are greatly reduced. Because of thegreat flexibility of design and fabrication, many functions can also beincorporated into it easily.

6.19. Pump and Valve Structures

The invention provides systems for fabricating and operatingmicrofabricated structures such as on/off valves, switching valves andpumps made out of various layers of elastomer bonded together. Thesestructures are suitable for controlling and fluid movement in thedescribed devices, e.g. flow control in the fluid or treatment channelsand circulation in a target hybridization loop.

As described, the invention uses multilayer soft lithography to buildintegrated (i.e.: monolithic) microfabricated elastomeric structures.Layers of soft elastomeric materials are bound together, resulting inbiocompatible devices that are reduced by more than two orders ofmagnitude in size, compared to conventional silicon-based devices. Thepreferred elastomeric material is a two-component addition cure materialin which one layer (e.g. a bottom layer) has an excess of one component,while another adjacent layer has an excess of another component. In anexemplary embodiment, the elastomer used is silicone rubber. Two layersof elastomer are cured separately. Each layer is separately cured beforethe top layer is positioned on the bottom layer. The two layers are thenre-cured to bond the layers together. Each layer preferably has anexcess of one of the two components, such that reactive molecules remainat the interface between the layers. The top layer is assembled on topof the bottom layer and heated. The two layers bond irreversibly suchthat the strength of the interface approximates or equals the strengthof the bulk elastomer. This creates a monolithic three-dimensionalpatterned structure composed entirely of two layers of bonded togetherelastomer. When the layers are composed of the same material, interlayeradhesion failures and thermal stress problems are avoided. Additionallayers may be added by repeating the process, wherein new layers, eachhaving a layer of opposite “polarity” are cured and bonded together.

Thus, in a preferred aspect, the various layers of elastomer are boundtogether in a heterogenous (A to B) bonding. Alternatively, a homogenous(A to A) bonding may be used in which all layers would be of the samechemistry. Thirdly, the respective elastomer layers may optionally beglued together by an adhesive instead.

Elastomeric layers may be created by spin coating an RTV mixture on amold at 2000 rpms for 30 seconds yielding a thickness of approximately40 microns. Layers may be separately baked or cured at about 80° C. for1.5 hours. One elastomeric layer may be bonded onto another by baking atabout 80° C. for about 1.5 hours. Micromachined molds may be patternedwith a photoresist on silicon wafers. In an exemplary aspect, a ShipleySJR 5740 photoresist was spun at 2000 rpms patterned with a highresolution transparency film as a mask and then developed yielding aninverse channel of approximately 10 microns in height. When baked at2000° C. for about 30 minutes, the photoresist reflows and the inversechannels become rounded. In preferred aspects, the molds may be treatedwith trimethylchlorosilane (TMCS) vapor for about a minute before eachuse in order to prevent adhesion of silicone rubber.

In another preferred aspect, a first photoresist layer is deposited ontop of a first elastomeric layer. The first photoresist layer is thenpatterned to leave a line or pattern of lines of photoresist on the topsurface of the first elastomeric layer. Another layer of elastomer isthen added and cured, encapsulating the line or pattern of lines ofphotoresist. A second photoresist layer is added and patterned, andanother layer of elastomer added and cured, leaving line and patterns oflines of photoresist encapsulated in a monolithic elastomer structure.Thereafter, the photoresist is removed leaving flow channel(s) andcontrol line(s) in the spaces which had been occupied by thephotoresist. Tetrabutylaminonium is one photoresist etchant that iscompatible with a preferred RTV 615 elastomer. An advantage ofpatterning moderate sized features (10 microns) using a photoresistmethod is that a high resolution transparency film can be used as acontact mask. This allows a single researcher to design, print, patternthe mold, and create a new set of cast elastomer devices, typically allwithin 24 hours.

A preferred elastomeric material is GE RTV 615 elastomer or a siliconerubber that is transparent to visible light, making multilayer opticaltrains possible. This allows optical interrogation of various channelsor chambers in the microfluidic device. In addition, GE RTV 615elastomer is biocompatible. Being soft, closed valves form a good sealeven if there are small particulates in the flow channel. Siliconerubber is also biocompatible and inexpensive, especially when comparedwith a crystal silicon.

The systems of the invention may be fabricated from a wide variety ofelastomers, such as the described silicon rubber and RTV 615. However,other suitable elastomeric materials may also be used. GE RTV 615(formulation) is a vinyl silane crosslinked (type) silicone elastomer(family). The invention is not limited to this formulation, type or eventhis family of polymer; rather, nearly any elastomeric polymer issuitable. An important requirement for the preferred method offabrication of the present microvalves is the ability to bond multiplelayers of elastomers together. In the case of multilayer softlithography, layers of elastomer are cured separately and then bondedtogether. This scheme requires that cured layers possess sufficientreactivity to bond together. Either the layers may be of the same type,and are capable of bonding to themselves (A to A), or they may be of twodifferent types, and are capable of bonding to each other (A to B).(Another possibility is to use an adhesive between layers.)

Given the tremendous diversity of polymer chemistries, precursors,synthetic methods, reaction conditions, and potential additives, thereare a huge number of possible elastomer systems that could be used tomake monolithic elastomeric microvalves and pumps. Variations in thematerials used will most likely be driven by the need for particularmaterial properties, i.e. solvent resistance, stiffness, gaspermeability, or temperature stability. There are many, many types ofelastomeric polymers. A brief description of the most common classes ofelastomers is presented here, with the intent of showing that even withrelatively “standard” polymers, many possibilities for bonding exist.Common elastomeric polymers include polyisoprene, polybutadiene,polychloroprene, polyisobutylene, poly(styrene-butadiene-styrene), thepolyurethanes, and silicones. See e.g., Ser. No. 60/186,856 filed Mar.3, 2000.

In addition to the use of “simple” or “pure” polymers, crosslinkingagents may be added. Some agents (like the monomers bearing pendantdouble bonds for vulcanization) are suitable for allowing homogeneous (Ato A) multilayer soft lithography or photoresist encapsulation;complementary agents (i.e. one monomer bearing a pendant double bond,and another bearing a pendant Si—H group) are suitable for heterogeneous(A to B) multilayer soft lithography.

Materials such as chlorosilanes such as methyl-, ethyl-, andphenylsilanes, for example polydimethylsilooxane (PDMS) such as DowChemical Copr. Sylgard 182, 184 or 186, or alipathic urethanediacrylates such as (but not limited to) Ebecryl 270 or In 245 from UBCChemical may also be used. Elastomers may also be “doped” withuncrosslinkable polymer chains of the same class. For instance RTV 615may be diluted with GE SF96-50 Silicone Fluid. This serves to reduce theviscosity of the uncured elastomer and reduces the Young's modulus ofthe cured elastomer. Essentially, the crosslink-capable polymer chainsare spread further apart by the addition of “Inert” polymer chains, sothis is called “dilution”. RTV 615 cures at up to 90% dilution, with adramatic reduction in Young's modulus.

The described monolithic elastomeric structures valves and pumps can beactuated at very high speeds. For example, the present inventors haveachieved a response time for a valve with aqueous solution therein onthe order of one millisecond, such that the valve opens and closes atspeeds approaching 100 Hz. The small size of these pumps and valvesmakes them fast and their softness contributes to making them durable.Moreover, as they close linearly with differential applied pressure,this allows fluid metering and valve closing in spite of high backpressures.

In various aspects of the invention, a plurality of first flow channelspass through the elastomeric structure with a second flow channel, alsoreferred to as an air channel or control line, extending across andabove a first flow channel. In this aspect of the invention, a thinmembrane of elastomer separates the first and second flow channels.Movement of this membrane (due to the second flow channel beingpressurized) will cut off flow passing through the lower flow channel.Typically, this movement is downward from a the interface with topcontrol layer into an closing an underlying first flow channel.

A plurality of individually addressable valves can be formed andconnected together in an elastomeric structure, and are then activatedin sequence such that peristaltic pumping is achieved. In other optionalpreferred aspects, magnetic or conductive materials can be added to makelayers of the elastomer magnetic or electrically conducting, thusenabling the creation of elastomeric electromagnetic devices.

In preferred aspects, channels of the invention have width-to-depthratios of about 10:1. In an exemplary aspect, fluid and/or air channelshave widths of about 1 to 1000 microns, and more preferably 10-200microns and most preferably 50-100 microns. Preferred depths are about 1to 100 microns, and more preferably 2-10 microns, and most preferably 5to 10 microns.

In preferred aspects, an elastomeric layer has a thickness of about 2 to2000 microns, and more preferably 5 to 50 microns, and most preferably40 microns. Elastomeric layers may be cast thick for mechanicalstability. In an exemplary embodiment, one or more layers is 50 micronsto several centimeters thick, and more preferably approximately 4 mmthick. Membrane separating fluid and air channels has a typicalthickness of about 30 nm. In one embodiment the thickness of oneelastomeric layer (e.g. at top or control layer) is about 10 times thethickness of an adjacent layer (e.g. a fluid or bottom layer.

A typical RTV valve of the invention is 100 μm×100 μm×10, connected toan off-chip pneumatic valve by a 10-cm-long air tube. In one example,the pressure applied on the control line is 100 kPa, which issubstantially higher than the approximately 40 kPa required to close thevalve. Thus, when closing, the valve in this example is pushed closedwith a pressure 60 kPa greater than required. When opening, however, thevalve is driven back to its rest position only by its own spring force,which is less than or equal to about 40 kPa). A signal to open or closethe valve is effected by changing the pressure accordingly. In thisexample there is a lag between the control signal and the controlpressure response, due to the limitations of the miniature valve used tocontrol the pressure. To accommodate this lag, these exemplary valvesrun comfortably at 75 Hz when filled with aqueous solution. If one usedanother actuation method which did not have an opening and closing lag,this valve would run at about 375 Hz. Note also that the spring constantcan be adjusted by changing the membrane thickness; this allowsoptimization for either fast opening or fast closing.

The flow channels of the present invention may optionally be designedwith different cross sectional sizes and shapes, offering differentadvantages, depending upon their desired application. For example, thecross sectional shape of a lower fluid channel may have a curved uppersurface, either along its entire length or in the region disposed underan upper air channel or cross channel. In certain embodiments a curvedupper surface facilitates valve sealing. In an alternate aspect, thebottom of a fluid channel is rounded such that its curved surface mateswith the curved upper wall upon valve closure.

6.20. Arrayable Rotary Mixer

This example describes a preferred embodiment of a microfluidic device.In particular, the example describes a microfluidic device that has aplurality (i.e., at least two) of hybridization or target loops, withthe fluid flow in each hybridization or target loop being driven by aperistaltic pump. Each target loop in the device may be connected to aseparate input well and/or a separate output well (e.g., via separateinput and/or output channels) to permit simultaneous loading andanalysis of several samples in the microfluidic device.

Alternatively, several of the hybridization or target loops may beconnected to common input and/or output wells; for example, by commoninput and/or output channels that branch into several branch channels,with each branch channel connecting to a different hybridization ortarget loop in the microfluidic device. A single sample (e.g., a singlebiological sample, such as a nucleic acid sample from a single patient)may then be loaded into the microfluidic device via an input well, and aplurality of different assays may be simultaneously performed on thesample. In particular, each hybridization or target loop may perform aseparate biological assay, such as detecting the presence of a differentnucleic acid (e.g., the presence of a different gene or a differentmutation of a particular gene or genes).

In particularly preferred embodiment, such a microfluidic device may becombined with standard microtiter plates, for example standard 96-wellor 384-well microtiter plates, that are widely used in most biologicallaboratories. Such a device will then be compatible with most existingloading and/or reading systems of loading and analyzing biologicalsamples.

FIGS. 21A and 21B illustrate two preferred embodiments of singlehybridization or target loops and their associated peristaltic pumpswhich can be arrayed to form a microfluidic device of this example. Thedevice comprises a fluid or “treatment” layer (bottom layer) containingthe loop channel (2101) and an air channel or “control” layer (toplayer) having a plurality of air channels (2102). The plurality of airchannels in the air channel layer form microvalves (2103) where each airchannel intersects the loop channel.

In this example, the fluid channels are about 300 μm wide and about 30μm deep. Suitable fluid channels are identical to those describedelsewhere in this specification for microfluidic devices (see, e.g.,Section 6.14). Typical fluid channel dimensions include those rangingfrom about 5-1000 μm wide and from about 1-50 μm deep, more preferablyabout 10-200 μm wide and about 2-30 μm deep. A preferred depth is about10 μm.

The exemplary loop channels shown in FIGS. 21A and 21B form loops havingthe geometry described in Section 6.18.2, supra, and illustrated in FIG.20. However, the loop channels may form any shape of loop, including anyof the particular shapes described in this specification. For example,the hybridization or target loops may be circular loops (as shown inFIG. 13B) or they may be rectangular (e.g., square) loops (as shown inFIG. 14). Loops having a geometry as described in Section 6.18.2 (e.g.,the geometry illustrated in FIG. 20) are preferred.

The air channels (2102) in this example are about 100 to 300 μm wide andabout 30 μm deep. Suitable air channel dimensions include those rangingfrom about 10-1000 μm wide (more preferably about 50-200 μm wide) andabout 2-50 μm deep (more preferably about 10-50 μm deep). A preferredparticularly depth is about 20 μm. In this embodiment, the air channelsare preferably parallel (or antiparallel) and do not intersect. The airchannels preferably run across the entire target or hybridization loopformed by the loop channel (i.e., they completely traverse the areaencompassed by the loop).

In preferred embodiments, the loop channels in a microfluidic devicehave at least one, and more preferably a plurality of channel supports(2104). In general, a channel support (2104) is located in the loopchannel at a point where an air channel (2102) intersects the loopchannel. The channel support supports the membrane above the fluidchannel (i.e., between the air channel layer and the fluid channellayer) without blocking the fluid channel; e.g. . . . , fluid can flowsthrough the fluid channel around the channel support. Where anair-channel intersects a fluid channel at a point (2103) not having achannel support, the membrane between the air and fluid channels is notpushed down by application of a sufficient air pressure in the airchannel. Thus, application of the air pressure causes the air channel tofunction as a microvalve and restrict or close the fluid channel atpoint 2103. However, where an air channel intersects a fluid channel ata point having a channel support 2104, the channel support prevents theair channel from restricting or closing the fluid channel. Thus, thesepoints do not function as microvalves.

Microchannels having channel supports may be readily obtained, e.g.,using any of the microfabrication techniques described supra. Forexample, in preferred embodiments individual layers for a microfluidicdevice are prepared from fluid molds fabricated on silicon wafers usingphotolithography (see, e.g., Sections 6.9 and 6.13.2, supra). In suchembodiments, standard micromachining techniques may be used, e.g., tocreate a negative master mold out of a silicon wafer. The mold may havea positive channel contour (see, e.g., FIG. 8) with a “hole” in itscenter. Curing a silicone elastomer (e.g., RTV 615) over such a moldthereby creates a microfluidic layer with a negative channel having achannel support therein. In other embodiments, microchannel supports maybe manufactured in the control layer (e.g., in air channels) instead ofor in addition to the fluid channels. Alternatively, functionallyequivalent channel supports may be obtained by decreasing the width of afluid channel, an air channel or both at a point where the fluid channeland the air channel intersect.

The microchannel supports may be any shape. However, circular or squaresupport shapes are preferred. Preferably, the supports are about 2-100μm wide, and are more preferably about 5-30 μm wide. The skilled artisanwill further appreciate that the invention may include embodiments wheremultiple supports are located at a channel intersection (e.g., two ormore, three or more, four or more, or five or more supports). A skilledartisan can readily determine appropriate spacing between multiplesupports according to the channel depth and actuation pressure (e.g.,the force per unit area) in the air channel of a particular microfluidicdevice. However, typical support spacing is between about 5-500 μm, andis more preferably between 10-100 μm.

FIG. 21A shows one embodiment of an arrayble loop channel which istraverse by three parallel air channels (2102). Although each channeltypically intersects the loop channel at two or more points, only oneintersection point is not blocked or occluded by a channel supports.Thus, each air channel intersects a loop channel at only one microvalve(2103).

An alternative embodiments of an arrayble loop channel is illustrated inFIG. 21B. The loop channel of this particular embodiment is traversed byfour parallel air channels (2102) which have a wider width (e.g., about20-1000 μm) at points where they intersect the loop channel to formmicrovalves (2103). However, the air channels have narrower widths(e.g., about 5-100 μm) at other points where they intersect the loopchannel (2101), thereby forming structures that function as “channelsupports” (2104) and prevent restriction or closing of the fluid channelwhen pressure is applied to the air channels. In the particularembodiment illustrated in FIG. 21B, the air channels may also traversethe loop channel along channel walls and/or between parallel andantiparallel microchannels (2105 and 2106). Preferably, the air channelsare narrower (e.g., between about 5-100 μm wide) and, more preferably,are no wider than the separation distance between the parallel andantiparalell channels.

Although any number of air channels may traverse or intersect anarrayble loop channel, there are preferably at least three air channelstraverse a given loop channel to form at least three microvalves. Theremay, however, be 4, 5, 6, 7 or more air channels intersecting a loopchannel to form 4, 5, 6, 7 or more microvalves.

FIGS. 22A and 22B illustrate an exemplary microfluidic device thatcomprises arrays of target or hybridization loops 2201. FIG. 22A showsone preferred embodiment of such a device that comprises an array of 96target or hybridization loops which are compatible with the wells of astandard 96-well microtiter plate. FIG. 22B is an exemplary “4-cell”microfluidic device (i.e., a microfluidic device comprising four targetor hybridization loops).

As with other microfluidic devices of this invention, these microfluidicdevices comprise a fluid channel or “treatment” layer (bottom layer)that contains the microfluidic channels (including the loop channels),and an air channel layer or “control” (top layer). The air channel layerin the device comprises a plurality of parallel air channels (2202) thateach traverse a plurality of the loop channels (2201). As each airchannel 2202 traverse a loop channel 2201 it preferably intersects theloop channel at one point that does not have a channel support, therebyforming a microvalve (2203). The fluid layer in such microfluidicdevices preferably comprises one or more additional fluid channels(2204), such as an inlet channel and/or an outlet channel feeding intoeach loop channel of the device. The inlet and/or outlet channels may bein fluid connection with a single inlet well or outlet well (e.g., forfeeding a single sample into each of the target or hybridization loops).Alternatively, the inlet channel, the outlet channel or both the inletand the outlet channels for each target or hybridization loop may be influid connection with a separate sample inlet or with a separate sampleoutlet (e.g., so that a separate sample may be loaded and/or analyzed ineach target or hybridization loop).

As with other embodiments of this invention, target molecules (e.g., DNAhybridization probes or antibody probes) are preferably laid down alongthe loop channels, e.g., on a glass substrate (2205) which lies beneaththe fluid channel layer. Preferably, different target molecules (e.g.,DNA hybridization probes having different nucleic acid sequences) arelaid down along each loop channel. For example, as with the embodimentsdescribed in Section 6.13, supra, the channels of the microfluidicdevices are preferably exposed to a common face of a chip. Themultiplayer assembly (for example an RTV assembly, as described inSection 6.13 supra) may be aligned and bonded to a substrate (e.g., aglass or other transparent substrate) so that a hermetic seal formsbetween the fluid channels and the substrate. The substrate ispreferably patterned in advance with one or more different targetmolecules (e.g., one or more sets of DNA or antibody probes) atpositions corresponding to the loop channels (2201).

A variety of additional microfluidic structures and functions may alsobe incorporated in such microfluidic devices, including any of the othermicrofluidic structures and functions described herein. For example,FIG. 22B illustrates a microfluidic device having additional channelswhich may be used, e.g., for the delivery of common samples, reagents,buffers or other chemicals to the target or hybridization loops. Otherchannels may be included, e.g., for the removal of sample, reagents,buffers, etc. from the target or hybridization loops (e.g., wastechannels). Microvalves and/or channel supports may also bemicrofabricated and/or incorporated into such other channels.

In particularly preferred embodiments, a microfluidic device such as theone shown in FIG. 22A is readily designed so that the spacing anddimensions of the target loop channels correspond to and/or arecompatible with the wells of a standard microtiter plates (e.g., astandard 96-well or 384-well plate). Each well of a microtiter plate canbe loaded with target molecules (preferably different target molecules)and the fluid channel layer for the device are laid over and,optionally, bonded to the microtiter plate.

Using such microfluidic devices, the amount of sample volume needed foranalysis is greatly reduced. For example, in preferred embodimentssample volumes between about 1-500 μl, and more preferably between about1-100 μl, can be loaded into a sample well of the microfluidic device.In other embodiments, sample volumes of less than 1 μl can be used. Forexample, in preferred embodiments the sample volume may be from about1-1000 nl, more preferably between about 1-500 nl, still more preferablybetween about 1-100 nl, and even more preferably from about 1 nl toabout 10 nl, about 20 nl, or about 50 nl. In particularly preferredembodiments, the sample volume loaded into a microfluidic device may beas low as 500 pl or less (e.g., as low as 200 pl, 100 pl, 50 pl, orless). In addition, because the pumps and valves of microfluidic devicesin this invention have negligible volumes, they are able to transportfluids at up to a few femtoliters per second, and more preferably at upto a few nanoliters per second.

Thus, smaller volumes of sample may be analyzed in parallel in largenumbers of microtiter wells. Accordingly, in one embodiment, amicrofluidic device of the invention comprises an array of 96-loopchannels and/or are compatible with a 96-well plate. In anotherpreferred embodiment, a microfluidic device of the invention comprisesan array of 384 loop channels and/or is compatible with a 384-wellplate. In other embodiments, the devices comprise even more numbers ofloop channels and may be used with even larger microtiter plates. Forexample, in another preferred embodiment the device comprises 1536 loopchannels and/or is compatible with a 1536-well plate.

6.21. Rotary Puming and Inline Mixing

This example presents experimental data that were obtained using amicrofluidic device having a chip architecture as described in Section6.13, supra. Specifically, the microfluidic device used in this exampleis one having a rotary loop and driven by a peristaltic pump. The datapresented here confirm that fluids may be rapidly mixed using such adevice. As such, the devices of this invention offer superior resultscompared to similar devices that are currently available.

6.21.1. Device Design and Fabrication

The microfluidic device used in this example comprised the samearchitecture and was identical to the device illustrated in FIG. 14 anddescribed in Section 6.13, above. FIG. 17 is a schematic diagram of thedevice used in these experiments. A photograph of that device isprovided in FIG. 13. Briefly, the device consists of two layers: a firstor “bottom” layer comprising fluid channels, and a second or “top” layercomprising pneumatic actuation channels. The bottom fluid layer also hastwo sample inputs (labeled in₁ and in₂ in FIG. 17), a mixing T-junction,a central circulation loop (also referred to here as a “rotary loop”),and an output channel. The loop diameter is 2.4 mm, while the channeldimensions in this exemplary device are 100 μm wide by 10 μm deep. Thetop layer has several stand alone actuation channels, which can bepressured or vented to atmosphere. Any intersection of a top air channelwith a bottom fluid channel forms a microvalve. The valve is closed whenthe air channel is pressurized, and is released otherwise. Several inletand outlet valves were built in to control the flow of each individualfluid component. When a series of on/off actuation sequencers (forexample, 001, 011, 010, 110, 100, 101) are applied to the air lines, thefluid can be peristaltically pumped through the loop in a chosendirection, which can be either clockwise or counterclockwise (see, also,Section 6.14 for detailed explanation of such actuation sequences). Thehigher the actuation frequency, the faster the fluid flows or “rotates”through the loop.

The device was fabricated using multilayer soft lithography techniquesthat are described here in Section 6.13.2, supra. See, also, the recentpublication of Unger et al (76) and by Chou (88). See, also, U.S.provisional patent application Ser. No. 60/249,362 filed on Nov. 16,2000: Briefly, mother molds for top and bottom layer were firstfabricated on silicon wafers by photolithography with photoresist(Shipley SJR 5740). Channel heights were controlled precisely by thespin coating rate. After UV exposure and development, photoresistchannels were formed. Heat reflow process and protection treatment wereapplied (10). Then, mixed two-part silicone elastomer (GE RTV 615) wasspun onto the bottom mold and poured onto the top mold, respectively.Again, spin coating was used to control the thickness of the bottompolymeric fluid layer. After baking in the over at 80° C. for 25minutes, the partially cured top layer was peeled off from its mold,aligned and assembled with the bottom layer. A 1.5-hour final bake at80° C. was applied to bind these two layers irreversibly. Once peeledoff from the bottom silicon mother mold, this RTV device was sealedhermetically to a glass cover slip.

6.21.2. Rotary Pumping and Inline Mixing

Fluid in a microfabricated rotary loop of this invention can be pumpedperistaltically when a proper actuation sequence is applied, asdescribed supra. To demonstrate such pumping, the central loop wasloaded with 2.5 μm diameter fluorescent beads. Valves regulating flowinto or out of the loop (i.e., the inlet and outlet valves) were closed,and the peristaltic pump was turned on by applying an appropriateactuation sequence as described above. Clear circulation of the beadsaround the loop could be visualized without any net flux into or out ofthe loop. By simply controlling the frequency of the actuation sequenceapplied with the peristaltic pump, the speed or flow-rate of particlesaround the loop could be adjusted. Similarly, the rotation directioncould be changed simply by reversing the actuation sequence.

In a microfabricated device of this invention, sample may be loaded intoa microfabricated rotary loop (e.g., through one or more inlet channels)and quickly mixed using the rotary pump. Without being limited to anyparticular theory or mechanism of action, it is understood that as twoor more different fluids rotate in the same loop, fluid at the center ofthe loop channel flows faster than fluid located at the channel's edge.As a result, it is expected that the interface between the two fluidswill continue to stretch as the fluids circulate, until each fluidbecomes a long, thin stream that wraps around the other fluid.Components of the different fluids can then cross the interface betweenthe fluids by diffusion, allowing the fluids to quickly and efficientlymix. By contrast; other microfluidic devices that are currentlyavailable typically inject different fluids into a chamber, and wait forthem to mix by slow diffusion. Thus, microfabricated devices having arotary loop of this invention offer particular advantages over existingdevices.

To demonstrate the device's ability to effectively and efficiently mixdifferent fluid components, experiments were performed in which themicrofluidic device schematically illustrated in FIG. 17 was used for“fixed-volume mixing” and “continuous-flow mixing.” The results fromthese two experiments, which are similar to the experiments described,above, in Section 6.15, are presented herebelow.

Fixed-volume Mixing. The term “fixed-volume mixing”, as used here,refers to instances in which two different solutions (for example, of asample and a reagent) of fixed volume are mixed completely before beingdirected to another stage of processing. To demonstrate the ability of amicrofabricated rotary loop to mix fluids under these circumstances, asolution of the fluorescent dye FITC was loaded into one of the device'sinput channels (in₁), and a solution of 1 μm fluorescent beads wasloaded into the other input channel (in₂). Because of laminar flow andthe slow diffusion rate, these two streams of fluid did not mix as theyentered the T-junction of the device. Instead, each fluid remainedconfined to one half of the flow channel, and the two fluid flowed sideby side into the central loop. Upon entering the central loop, the twofluids split into their two distinct parts and met again at the bottomof the loop, as shown in FIG. 18A. Thus, one half of the central loopwas filled with a solution containing the FITC dye, while the other halffilled with a solution containing the fluorescent beads.

Once the central loop filled with fluid as described above, theperistaltic pump was activated with the appropriate actuation sequenceand at a frequency of 30 Hz. After only 30 seconds of pumping, both thefluorescent dye and the beads were uniformly distributed throughout thewhole central loop, as shown in FIG. 18B. Thus, a microfabricated rotaryloop of this invention can be used to rapidly mix two or more differentreagents, and can be particularly useful when one or more of thedifferent components have slow diffusion constants; for example, with aDNA sample to be mixed with hybridization beads or cells to be mixedwith plasmid-lipid complexes.

Continuous-flow Mixing. The term “continuous-flow mixing” is used hereto describe mixing that takes place as two or more fluids continuouslyflow down a micro-channel. Thus, unlike the fixed-volume mixingexperiments described above, continuous-flow mixing refers to asituation where inlet and outlet channels to a rotary loop are notclosed so that solution flows continuously through the loop. A secondexperiment was performed to demonstrate the ability of a microfabricatedloop to mix such a continuous flow of fluids. This second experiment wasidentical to the fixed-volume mixing experiment described above, exceptthat the inlet and outlet channels were left open so that the bead anddye solutions flowed continuously through the loop at a flow rate ofabout 2 mm/s.

As in the fixed volume mixing experiment (supra), the two fluids did notmix as they entered the T-junction and, upon entering the loop, splitinto their two distinct parts and met again at the bottom of the loop(FIG. 23A). When the rotary pump was activated, complete dye mixing wasobserved (FIG. 23B). Partial mixing of the extremely slow 1 μm beadsalso occurred so that about one quarter (i.e., approximately 25%) of thebeads were sent to the other side of the central loop after passingthrough the continuous-flow rotary mixer only once.

To quantify the level of mixing, 119 video frames were analyzed to countbeads within a window channel, and to note their location. These dataare shown in FIGS. 23C and 23E for experiments where the rotary pump wasinactive and mixing occurred by diffusion, and in FIGS. 23D and 23F forexperiments where the rotary pump was activated. FIGS. 23C and 23D showa plot indicating the location where fluorescent beads were observedacross the channel width (the horizontal or “X” axis) and along a shortsection of the outlet channel length (the vertical or “Y” axis) in theanalyzed frames. FIGS. 23E and 23F plot the accumulated distribution offluorescent beads observed across the channel width (“X”).

The quality of continuous-flow mixing in such a rotary loop dependsstrongly on the ratio of the overall flow rate and the pump rate. Thus,the mixing quality can be optimized for a particular application byeither lowering the overall flow rate with which fluid is introducedinto the loop, by widening and/or lengthening the rotary mixer loop, orby increasing the rotary mixing speed (e.g., by increasing the actuatingsequence for the peristaltic pump).

Estimating rotary flow mixing times. Without being limited to anyparticular theory or mechanism of interaction, a simple model ispresented to estimate the mixing effect of the rotary flow in amicrofluidic device of this invention. In particular, the model providesformulas that are useful, e.g., to estimate the time required foreffective mixing using a microfluidic device of the invention withactive rotamer pumping (for example, in the fixed volume and/orcontinuous volume mixing experiments described above).

FIG. 24 provides a simplified schematic for an exemplary microfabricatedcentral loop. The microfabricated channel forming the loop has a radiusr₀, and forms a circular loop of radius R.

As described above, when two separate fluids (referred to here as fluidsA and B, respectively) enter the loop through a microfabricated inletchannel, they are not mixed. Instead, each fluid is confined to one halfof the flow channel so that the two fluids flow side by side into theloop. Upon entering the loop the two fluids separate. One fluid (e.g.,fluid A) travels counterclockwise through the loop, while the otherfluid (e.g., fluid B) travels clockwise through the loop until the twofluids meet again at the bottom of the loop (i.e., opposite the inletchannel). Thus, in the simple model described here fluid. A initiallyfills the loop over an angular range of 0 to π radians (i.e. 0 to 180°).Fluid B initially fills the loop over an angular range of π to 2πradians (i.e., from 180° to) 360°.

As fluid is pumped through the loop channel, it experiences adifferential velocity. In particular, fluid near the channel's edgeexperiences a dragging force from the channel wall, and therefore flowsat a slower speed. Fluid near the center of the loop channel experiencesless drag or resistance, and therefore travels at a faster speed.Accordingly, during active pumping the flow velocity v at a distance rfrom the channel center is estimated in the present model by

$\begin{matrix}{v = {v_{0}( {1 - ( \frac{r}{r_{0}} )^{2}} )}} & ( {{Eq}.\mspace{14mu} 6.21.A} )\end{matrix}$

where v₀ is the maximum fluid velocity (i.e., the velocity at r=0, thecenter of the fluid channel). Similarly, the angular velocity of fluidflow through the loop ω=v/R is estimated by the formula

$\begin{matrix}{\omega = {\omega_{0}( {1 - ( \frac{r}{r_{0}} )^{2}} )}} & ( {{Eq}.\mspace{14mu} 6.21.B} )\end{matrix}$

where ω₀=v₀/R (i.e., the angular velocity at r=0, the center of thechannel As fluid is pumped through the loop for a time t, the front orboundary between fluids A and B travels through an angular distancegiven by ωt·mod(2π).

From Equation 6.21B, above, it is easy to see that when the fluid centermoves through an angular distance ω₀t=π, the center of the front betweenfluids A and B will have traversed one-half turn around the loop.However, that part of the front adjacent to the channel wall will nothave moved at all. Thus, at time t=π/ω₀, the greatest distance overwhich fluids A and B must diffuse to mix is simply equal to the channelradius, r₀. From this logic, it can be seen that once the front at adistance r from the channel center has traversed one-half turn aroundthe loop, the greatest diffusion distance l is simply the distance tothe channel wall; i.e., l=r₀−r. The diffusion distance may therefore beestimated by solving Equation 6.21.C, below, for r₀−r.

$\begin{matrix}{{{\omega \; t} = {{{\omega_{0}( {1 - ( \frac{r}{r_{0}} )^{2}} )}t} = \pi}},} & ( {{Eq}.\mspace{14mu} 6.21.C} )\end{matrix}$

This equation has an approximate solution given by:

$\begin{matrix}{l = {{r_{0} - r} \approx \frac{\pi \; r_{0}}{\omega_{0}t}}} & ( {{Eq}.\mspace{14mu} 6.21.D} )\end{matrix}$

which may be rewritten in the more convenient form provided in Equation6.21.E, below, where l₀=r₀, and k=ω₀/π=v₀/Rπ.

$\begin{matrix}{l = \frac{l_{0}}{kt}} & ( {{Eq}.\mspace{14mu} 6.21.E} )\end{matrix}$

In this latter expression, k is therefore a constant coefficient thatdepends upon the total distance around the loop (i.e., 2πR) and thepumping speed.

The time τ required for an object to diffuse across a distance l can beprovided by the equation

$\begin{matrix}{\tau = \frac{l^{2}}{2\; D}} & ( {{Eq}.\mspace{14mu} 6.21.F} )\end{matrix}$

where D is the object's diffusion constant. If rotary pumping occurs,however, the diffusion distance l is given by Equation 6.21.E, above.Thus, by substituting the expression for l given in Equation 6.21.E andsolving Equation 6.21.F for τ=t, the rotary mixing time for effectivelymixing two fluids A and B can be estimated as:

$\begin{matrix}{\tau = ( \frac{r_{0}^{2}}{2\; k^{2}D} )^{1/3}} & ( {{Eq}.\mspace{14mu} 6.21.G} )\end{matrix}$

In contrast, other devices typically mix fluids by passive diffusionacross the channel radius, r₀. However, such diffusion requires a timeperiod given by the equation:

$\begin{matrix}{\tau = \frac{r_{0}^{2}}{2\; D}} & ( {{Eq}.\mspace{14mu} 6.21.H} )\end{matrix}$

By comparing Equations 6.21.G and 6.21.H, supra, it is apparent thatmixing by rotamer pumping in a microfluidic device of this inventiongreatly improves upon other devices that are presently available. Inparticular, the required mixing time in a rotamer loop of this inventionis much less sensitive to both the channel width r₀ and the diffusionconstant D. For large objects (e.g., proteins, beads, virions, cells,etc.), the diffusion constant is typically between 10⁻⁵ and 10⁻⁸ cm²/s.Thus, whereas mixing of such objects by passive diffusion will typicallyrequire a time factor on the order of 10³, the time factor required formixing by active rotary pumping in a device of this invention reduced bytwo orders of magnitude, i.e., to a factor of only 10.

6.22. Surface Binding Assay Using a Multiparameter Chip

This example describes experiments that are similar to those describedin Sections 6.14 and 6.16, supra, and that illustrate the utility of amicrofabricated rotary loop in binding assays. Specifically, anexemplary assay is demonstrated in which avidin labeled beads bind tobiotin in the central loop of a microfabricated device. These data showthat using such a rotamer pump greatly shortens the time required forbinding to occur, thereby offering improved results over traditionaldevices where binding is drive by simple diffusion.

6.22.1. Surface Patterning of a Microfluidic Chip

The particular microfluidic device described, supra, in Section 6.2.1was also used in the binding assays described here. However, for thesebinding experiments the device was hermetically sealed to a glass coverslip that had been chemically patterned with biotin to form abinding-assay substrate. Soft lithographic techniques for chemicalpatterning are well known in the art and have been described, e.g., forantibody recognition in microfluidic network (79) and for microcontactprinting of self-assembled monlayers (89). For these experiments anelastomeric print head was designed having a radial pattern of fluidicchannels. The print head was attached to a glass cover slip whosesurface had been functionalized in advance with carboxyl groups. Abiotin-amine conjugate and coupling solution (comprising EDC andSulfo-NHS) were deposited in a well in the middle of the print head,which then filled the microfluidic spokes by capillary action. As aresult, the print head generated a crude pattern of biotin derivatized,radial “spokes” on the cover slip. More complicated patterns ofderivatized micronetworks can be readily designed by actively pumpingreagents through an elastomeric print head in arbitrary patterns.

After the biotin coupling reaction was complete, channels in the printhead were flushed with buffer, the print head was removed, and themicrofluidic device was attached to the cover slip as described inSection 6.21, supra. The resulting microfluidic device is shownschematically in FIG. 17, and included a central loop (105) with eight100 μm by 100 μm biotin spots or “pixels” (110) in the regions where thebiotin derivatized spokes intersected the loop channel. The total volumein the central rotary-pump reaction chamber was 7.54 nanoliters.

6.22.2. Surface Binding Assay

After a biotin surface patterned microfluidic device was prepared asdescribed, supra, 1 μm fluorescent beads (Molecular Probes) coated withNeutrAvidin (a derivative of avidin with less nonspecific binding) wereintroduced into the central loop from the input channel. Once the loopwas filed, input and output microvalves were closed and the peristalticrotary pump was activated at a cycling frequency of 10 Hz. Within fourminutes, more than 80% of the NeutrAvidin coated beads had bound to thebiotinylated spots, as shown in FIG. 19.

Control experiments were also performed in which the peristaltic rotarypump was not activated. Thus, in these control experiments binding tothe biotin spots was driven only by the diffusion of the NeutrAvidinlabeled beads. However, this is expected to be an extremely slow processdue to the slow diffusion constant of those beads (about 2.5×10⁻⁹cm²/s). Indeed, it took 4 hours before the fluorescent signal from thebeads was localized to the biotinylated spots. By this time, most beadswith about 50 μm of a biotinylated spot had bound thereto. Clearlytherefore, binding by diffusion alone is an extremely slow process andbeads on the other side of the loop from a detection spot would neverhave the chance to bind thereto, as is necessary for multiplexed surfacebinding assays. Thus, use of a rotary loop and pump of the presentinvention greatly improved the reaction kinetics, reducing the reactiontime by a factor of at least 60.

In summary, an active DNA diagnosis chip using multilayer softlithography is described. DNA fragments are patterned on a glasssubstrate with appropriate surface chemistry combinations and by usingelastomeric fluidic channels. A monolithic elastomeric diagnosis deviceis aligned and attached to the derivatized surface. DNA samples arebrought in, e.g. to fluid channels, are circulated several times in ahybridization loop containing probes bound to an aligned andcommunication surface, and sample fluid (e.g. with non-binding DNA) isthen expelled from an outlet, under valve control. Fluorescent dye, usedto determine, detect, measure or image the hybridization, is brought infrom the same or another input channel. Fast mixing can also be donewith the active pumping mechanism. Laminar flow and diffusionconsideration in this low-Reynolds number regime can be overcome easilyby the active pumping agitation.

While the invention has been described with reference to specificmethods and embodiments, it will be appreciated that variousmodifications may be made without departing from the invention.

7. BIBLIOGRAPHY AND REFERENCES CITED

The following Bibliography provides the complete citations to thereferences cited in the above text. The references are provided merelyto clarify the description of the present invention and citation of areference either in the below Bibliography or in the specification aboveis not an admission that any such reference is “prior art” to theinvention described herein. Each reference cited in this application,including the references listed in the below Bibliography and any otherreferences cited in the above specification, is incorporated herein, byreference, in its entirety and to the same extent as if each referenceswas incorporated by reference individually in the above specification.

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1-92. (canceled)
 93. A method for analyzing a cell, the methodcomprising: (a) introducing a fluid comprising the cell onto amicrofluidic device; (b) lysing the cell on the microfluidic device torelease cell material from the cell; and (c) circulating at least someof the cell material in a loop path on the microfluidic device, suchthat said cell material circulates multiple times though the loop path,the microfluidic device further comprising a peristaltic pump tocirculate cell material through the loop path, wherein said peristalticpump comprises at least three microvalves actuated in a coordinatedfashion to move the cell material through the loop path.
 94. A methodaccording to claim 93, further comprising analyzing the at least some ofthe lysed cell material.
 95. A method according to claim 93, wherein theloop path has a rectangular, a square, or a circular shape.
 96. A methodaccording to claim 93, wherein the loop path comprises a plurality ofinterconnected channels.
 97. A method according to claim 96, wherein theloop path comprises at least one pair of interconnected parallel andantiparallel channels.
 98. A method according to claim 93, wherein theloop path comprises at least one chamber.
 99. A method according toclaim 93, wherein the microfluidic device comprises a plurality of looppaths.
 100. A method according to claim 93, wherein the at least some ofthe cell material comprises a molecule selected from a group consistingof a nucleic acid molecule, a polypeptide molecule, and an antibodymolecule.
 101. A method according to claim 93, comprising performingpolymerase chain reaction amplification of a nucleic acid molecule inthe cell material.